High resolution radio frequency medical imaging and therapy system

ABSTRACT

A method for imaging includes directing a plurality of radio frequency (RF) beams toward a target organ from a plurality of angles. The RF beams include one or more first pairs of the RF beams, each first pair including two of the RF beams that impinge on the target organ from opposite directions. RF signals reflected from the target organ are received responsively to the RF beams, the RF signals including one or more second pairs of the RF signals engendered respectively by the one or more first pairs of the RF beams. Local tissue parameters at multiple points in the target organ are extracted by jointly processing the RF signals in each of the second pairs. Images of the target organ are produced using the extracted local tissue parameters. Other embodiments described herein include methods for passive imaging, motion vector analysis, ablation, local heating and application of electromagnetic pressure.

CROSS-REFERENCE TO RELATED APPLICATIONS

This application is continuation of U.S. patent application Ser. No.11/997,913 filed on Feb. 5, 2008, now U.S Pat. No. 7,529,398, which is a371 of PCT/IL2006/000896 filed on Aug. 3, 2006, which claims the benefitof U.S. Provisional Patent Application 60/707,064, filed Aug. 9, 2005,which is incorporated herein by reference.

FIELD OF THE INVENTION

The present invention relates generally to medical imaging and therapysystems, and particularly to methods and systems for high resolutionradio frequency (RF) imaging and therapy.

BACKGROUND OF THE INVENTION

Medical imaging methods and systems use a variety of imaging modalities.Each modality can be characterized by its typical spatial and temporalresolution. For example, the following table shows typically achievablespatial and temporal resolution values of several known imagingmodalities:

Spatial resolution Temporal resolution (per axis) (three-dimensionalModality [mm] frame refresh rate) [Hz] Ultrasound 1 20-30 Singlepositron emission 5 10 computerized tomography (SPECT) Positron emission3 10 tomography (PET) Computerized tomography 0.5-1 10 (CT) Magneticresonance 0.5-1 10 imaging (MRI)

Some methods and systems use radio frequency (RF) based imaging. Forexample, U.S. Pat. No. 6,490,471, whose disclosure is incorporatedherein by reference, describes a single-frequency three-dimensional(3-D) microwave tomographic device capable of imaging a full scalebiological object. The device includes code-division software, whichcooperates with a microwave patch system to enable superficial imagingof biological systems. A cluster of antennas and transceivers is used toprovide microwave tomography (MWT) and electrical impedance tomography(EIT) integrated in a single 3-D microwave system for examining thebiological object from a number of views in real-time.

PCT International Publication WO 03/003907, which is incorporated hereinby reference, describes a system for microwave imaging via space-timebeam-forming. Microwave signals are transmitted from multiple antennalocations into an individual to be examined. Backscattered microwavesignals are received at multiple antenna locations, to provide receivedsignals from the antennas. The received signals are processed in acomputer to remove the skin interface reflection component of the signalat each antenna. The corrected signal data is processed by abeam-former. The beam-former is scanned over a plurality of differentlocations in the individual by changing time shifts, filter weights andtime-gating of the beam-former process. The output power may bedisplayed as a function of scan location, with regions of large outputpower corresponding to significant microwave scatterers such asmalignant lesions.

U.S. Pat. No. 6,448,788, whose disclosure is incorporated herein byreference, describes a method and apparatus for microwave imaging of aninhomogeneous target, in particular of biological tissue. The methodcompensates for the interactions between active and non-active antennas.Measured electric field data is processed in magnitude and phase form,so that unwrapped phase information may be used directly in the imagereconstruction. Initial finite element measurements and calculations areused to determine the perimeter dimensions of the target being examined.

U.S. Pat. No. 6,253,100, whose disclosure is incorporated herein byreference, describes a method for imaging an object, such as a diseasedhuman heart or bone, in a non-transparent medium, such as the humanbody. The method involves placing an array of transmitters and receiversin operational association with the medium. The transmitters generate aharmonic or pulse primary electromagnetic (EM) field, which propagatesthrough the medium. The primary field interacts with the object toproduce a scattered field, which is recorded by the receivers. Thescattered EM field components measured by the receivers are applied asan artificial EM field to generate a backscattering EM field. Crosspower spectra of the primary and backscattering fields or crosscorrelation between these fields produce a numerical reconstruction ofan EM hologram. The desired properties of the medium, such asconductivity or dielectric permittivity, are then derived from thishologram.

SUMMARY OF THE INVENTION

There is therefore provided, in accordance with an embodiment of thepresent invention, a method for imaging, including:

directing a plurality of radio frequency (RF) beams toward a targetorgan from a respective plurality of angles, the plurality of the RFbeams including one or more first pairs of the RF beams, each first pairincluding two of the RF beams that impinge on the target organ fromopposite directions;

receiving RF signals reflected from the target organ responsively to theRF beams, the RF signals including one or more second pairs of the RFsignals engendered respectively by the one or more first pairs of the RFbeams;

extracting local tissue parameters at multiple points in the targetorgan by jointly processing the RF signals in each of the second pairs;and

producing images of the target organ using the extracted local tissueparameters.

In some embodiments, directing the plurality of the RF beams includesforming a respective plurality of effective antennas directed to thetarget organ from the plurality of the angles by selectively activatingsubsets of radiating elements selected from an antenna array including aplurality of the radiating elements. In another embodiment, the antennaarray includes a cylindrical array surrounding the target organ, and theRF beams are parallel, with an offset no greater than one degree, to abase of the cylinder and point toward a central axis of the cylinderfrom multiple azimuth angles and heights.

In yet another embodiment, directing the plurality of the RF beamsincludes mechanically scanning one or more antennas so as to transmitfrom the plurality of the angles. Additionally or alternatively,directing the plurality of the RF beams includes mechanically scanningthe target organ with respect to one or more antennas so as to cause theRF beams generated by the antennas to impinge on the target organ fromthe plurality of the angles.

In a disclosed embodiment, directing the plurality of the RF beamsincludes transmitting one or more wideband RF pulses in each of the RFbeams. In another embodiment, transmitting the one or more pulsesincludes transmitting a sequence of two or more wideband RF pulses andphase-encoding the sequence by assigning respective phases to the pulsesdepending on positions of the pulses in the sequence.

In an embodiment, receiving the RF signals includes sampling thereflected RF signals using multiple analog-to-digital (A/D) convertershaving incremental time offsets with respect to one another. Receivingthe RF signals may include enhancing a range resolution of the reflectedRF signals using multiple outputs of the multiple analog-to-digital(A/D) converters.

In another embodiment, receiving the RF signals includes applying atime-dependent gain control (TGC) function to the received RF signals.

In yet another embodiment, the local tissue parameters include at leastone parameter selected from a group of parameters consisting of a localattenuation coefficient, a local reflection coefficient, a local timedelay and a local tissue dielectric property.

In still another embodiment, receiving the RF signals includes measuringfor each of the second pairs first and second reflection intensityprofiles indicating intensities of the RF signals in the each of thesecond pairs as a function of time, and jointly processing the RFsignals includes comparing the first and second reflection intensityprofiles. Comparing the first and second reflection intensity profilesmay include identifying in the first reflection intensity profile firstreflection peaks reflected from respective tissue interfaces in a firstdirection, identifying in the second reflection intensity profile secondreflection peaks reflected from the respective tissue interfaces in asecond direction opposite to the first direction, and calculatingcorrected values of the local tissue parameters responsively todifferences between the first and second reflection peaks.

In some embodiments, calculating the local tissue parameters includescorrecting at least one artifact selected from a group consisting of alocal time delay and a local attenuation in the first and secondreflection intensity profiles.

In an embodiment, producing the images of the target organ includesreconstructing a three-dimensional (3-D) representation of the localtissue parameters by calculating accumulated contributions of thecorrected values of the local tissue parameters of the second pairs atthe multiple points in the target organ.

In another embodiment, at least some of the RF beams overlap oneanother, and reconstructing the 3-D representation includes improving aspatial resolution of the 3-D representation using the overlapping RFbeams.

Calculating the accumulated contributions may include calculating, foreach beam, iso-time surfaces defining loci of some of the multiplepoints in the target organ having a particular propagation delay withrespect to an antenna directing the beam.

In an embodiment, directing the RF beams, receiving the RF signals andextracting the tissue parameters include continually scanning the targetorgan using the RF beams, and producing the images of the target organincludes producing a sequence of 3-D images that display a variation ofthe extracted tissue parameters over time.

In another embodiment, the sequence of the 3-D images has a frame rategreater than fifty 3-D frames per second. In yet another embodiment, theframe rate is greater than or equal to one hundred 3-D frames persecond.

In still another embodiment, the method includes tracking a temporalvariation of a tissue region by measuring differences among respectivelocations of the tissue region in the sequence of 3-D images.

In an embodiment, the produced images of the target organ have a spatialresolution better than 2 mm. In another embodiment, the spatialresolution is better than 1 mm.

In yet another embodiment, producing the images of the target organincludes differentiating between first and second different tissue typesusing the extracted local tissue parameters. Additionally oralternatively, producing the images of the target organ includesidentifying a tissue type using the extracted local tissue parameters.Further additionally or alternatively, producing the images of thetarget organ includes measuring a local conductivity in at least some ofthe multiple points in the target organ using the extracted local tissueparameters.

In an embodiment, the target organ includes a heart. Additionally oralternatively, at least some of the RF beams penetrate a body containingthe target organ to a depth greater than 20 cm in order to image atleast some of the multiple points in the target organ. In someembodiments, the depth is greater than 30 cm.

In a disclosed embodiment, directing the RF beams includes configuring afirst subset of the RF beams to use a first polarization and a secondsubset of the RF beams to use a second polarization different from thefirst polarization, and extracting the local tissue parameters includescalculating first values of the local tissue parameters responsively tothe first subset of the beams and second values of the local tissueparameters responsively to the second subset of the beams.

In another embodiment, receiving the RF signals includes filtering thereceived RF signals to produce first and second partial bandwidth RFsignals, and extracting the local tissue parameters includes calculatingfirst values of the local tissue parameters responsively to the firstpartial bandwidth RF signal and second values of the local tissueparameters responsively to the second partial bandwidth RF signal.

In yet another embodiment, the method includes inserting a contrastagent affecting at least one of the local tissue parameters into thetarget organ.

In still another embodiment, extracting the local tissue parametersincludes estimating tissue motion velocities at the multiple points inthe target organ by measuring Doppler spectra of the RF signals in threeor more of the RF beams.

In some embodiments, the method includes applying RF ablation to anablation region in the target organ by focusing an ablating signal onthe ablation region using at least some of the RF beams. Focusing theablating signal on the ablation region may include directing theablating signal based on the produced images of the target organ. Inanother embodiment, the method includes locally heating a region of thetarget organ by focusing a heating RF signal on the region using atleast some of the RF beams. In yet another embodiment, the methodincludes applying an electromagnetic pressure to a region of the targetorgan by focusing an RF signal on the region using at least some of theRF beams.

There is additionally provided, in accordance with an embodiment of thepresent invention, a method for imaging, including:

configuring a set of antennas so as to define three or more axes ofdirectional reception of radio frequency (RF) signals from a targetorgan at respective different angles;

using the set of antennas, passively sensing the RF signals emittedalong the three or more axes due to a local electrical activity signalgenerated in the target organ; and

determining a location coordinate of the local electrical activitysignal based on the sensed RF signals.

In an embodiment, passively sensing the RF signals includes sampling theRF signals at a sampling rate higher than 1 GHz and integrating thesampled RF signals over a duration greater than or equal to 1microsecond. In another embodiment, determining the location coordinateof the local electrical activity signal includes periodicallydetermining the location coordinate and displaying a variation of thelocation coordinate over time. In yet another embodiment, determiningthe location coordinate includes filtering the sensed RF signals toproduce respective narrowband signals, and applying an interferometrycalculation to the narrowband signals.

There is further provided, in accordance with an embodiment of thepresent invention, a method for imaging, including:

directing a plurality of radio frequency (RF) beams toward a targetorgan from a respective plurality of antenna locations, the plurality ofthe RF beams including one or more first pairs of the RF beams, eachpair including two of the RF beams that impinge on the target organ fromopposite directions;

receiving RF signals reflected from the target organ responsively to theRF beams, the RF signals including one or more second pairs of the RFsignals engendered respectively by the one or more first pairs of the RFbeams;

compensating for local tissue artifacts in the RF signals by jointlyprocessing the RF signals in each of the second pairs; and

calculating three-dimensional (3-D) velocity vectors of multiple pointsin the target organ with respect to the antenna locations using the RFsignals after compensating for the local tissue artifacts.

In an embodiment, calculating the 3-D velocity vectors includesevaluating Doppler spectra of the RE signals with respect to the antennalocations for each of the multiple points, identifying dominant spectralcomponents in the Doppler spectra and associating three or more of thedominant spectral components in respective three or more of the Dopplerspectra to produce a 3-D velocity vector estimate.

In another embodiment, associating the three or more dominant spectralcomponents includes identifying and discarding false associationsbetween dominant spectral components by comparing the 3-D velocityvector estimate to at least one estimate selected from a group ofestimates consisting of previous 3-D velocity vector estimates and 3-Dvelocity vector estimates of adjacent points in the target organ.

There is also provided, in accordance with an embodiment of the presentinvention, an imaging system, including:

one or more antennas, which are arranged to direct a plurality of radiofrequency (RF) beams toward a target organ from a respective pluralityof angles, the plurality of the RF beams including one or more firstpairs of the RF beams, each first pair including two of the RF beamsthat impinge on the target organ from opposite directions;

a receiver, which is arranged to receive via the one or more antennas RFsignals reflected from the target organ responsively to the RF beams,the RE signals including one or more second pairs of the RF signalsengendered respectively by the one or more first pairs of the RF beams;and

a processor, which is arranged to extract local tissue parameters atmultiple points in the target organ by jointly processing the RF signalsin each of the second pairs and to produce images of the target organusing the extracted local tissue parameters.

There is additionally provided, in accordance with an embodiment of thepresent invention, an imaging system, including:

a set of antennas, which are configured to define three or more axes ofdirectional reception of radio frequency (RF) signals from a targetorgan at respective different angles;

a receiver, which is arranged to passively sense, using the set ofantennas, the RF signals emitted along the three or more axes due to alocal electrical activity signal generated in the target organ; and

a processor, which is arranged to determine a location coordinate of thelocal electrical activity signal based on the sensed RF signals.

There is further provided, in accordance with an embodiment of thepresent invention, an imaging system, including:

a set of antennas, which are arranged to direct a plurality of radiofrequency (RF) beams toward a target organ from a respective pluralityof antenna locations, the plurality of the RF beams including one ormore first pairs of the RF beams, each first pair including two of theRF beams that impinge on the target organ from opposite directions;

a receiver, which is arranged to receive via the set of antennas RFsignals reflected from the target organ responsively to the RF beams,the RF signals including one or more second pairs of the RF signalsengendered respectively by the one or more first pairs of the RF beams;and

a processor, which is arranged to compensate for local tissue artifactsin the RF signals by jointly processing the RF signals in each of thesecond pairs, and to calculate three-dimensional (3-D) velocity vectorsof multiple points in the target organ with respect to the antennalocations using the RF signals after compensating for the local tissueartifacts.

There is also provided, in accordance with an embodiment of the presentinvention, a computer software product for imaging, the productincluding a computer-readable medium, in which program instructions arestored, which instructions, when read by a computer, cause the computerto control one or more antennas to direct a plurality of radio frequency(RF) beams toward a target organ from a respective plurality of angles,the plurality of the RF beams including one or more first pairs of theRF beams, each first pair including two of the RF beams that impinge onthe target organ from opposite directions, to receive via the one ormore antennas RF signals reflected from the target organ responsively tothe RF beams, the RF signals including one or more second pairs of theRF signals engendered respectively by the one or more first pairs of theRF beams, to extract local tissue parameters at multiple points in thetarget organ by jointly processing the RF signals in each of the secondpairs and to produce images of the target organ using the extractedlocal tissue parameters.

There is additionally provided, in accordance with an embodiment of thepresent invention, a computer software product for imaging, the productincluding a computer-readable medium, in which program instructions arestored, which instructions, when read by a computer, cause the computerto configure a set of antennas to define three or more axes ofdirectional reception of radio frequency (RF) signals from a targetorgan at respective different angles, to passively sense, using the setof antennas, the RF signals emitted along the three or more axes due toa local electrical activity signal generated in the target organ, and todetermine a location coordinate of the local electrical activity signalbased on the sensed RF signals.

There is further provided, in accordance with an embodiment of thepresent invention, a computer software product for imaging, the productincluding a computer-readable medium, in which program instructions arestored, which instructions, when read by a computer, cause the computerto control a set of antennas to direct a plurality of radio frequency(RF) beams toward a target organ from a respective plurality of antennalocations, the plurality of the RF beams including one or more firstpairs of the RF beams, each first pair including two of the RF beamsthat impinge on the target organ from opposite directions, to receivevia the set of antennas RF signals reflected from the target organresponsively to the RF beams, the RF signals including one or moresecond pairs of the RF signals engendered respectively by the one ormore first pairs of the RF beams, to compensate for local tissueartifacts in the RF signals by jointly processing the RF signals in eachof the second pairs, and to calculate three-dimensional (3-D) velocityvectors of multiple points in the target organ with respect to theantenna locations using the RF signals after compensating for the localtissue artifacts.

There is additionally provided, in accordance with an embodiment of thepresent invention, a method for radio frequency (RF) ablation,including:

directing a plurality of RF beams toward a target organ from arespective plurality of angles, the plurality of the RF beams includingone or more first pairs of the RF beams, each first pair including twoof the RF beams that impinge on the target organ from oppositedirections;

receiving RF signals reflected from the target organ responsively to theRF beams, the RF signals including one or more second pairs of the RFsignals engendered respectively by the one or more first pairs of the RFbeams;

extracting local tissue parameters at multiple points in the targetorgan by jointly processing the RF signals in each of the second pairs;and

focusing an ablating signal on an ablation region in the target organusing multiple ablation beams based on the extracted local tissueparameters.

There is also provided, in accordance with an embodiment of the presentinvention, a radio frequency (RF) ablation system, including:

one or more antennas, which are arranged to direct a plurality of RFbeams toward a target organ from a respective plurality of angles, theplurality of the RF beams including one or more first pairs of the RFbeams, each first pair including two of the RF beams that impinge on thetarget organ from opposite directions;

a receiver, which is arranged to receive via the one or more antennas RFsignals reflected from the target organ responsively to the RF beams,the RF signals including one or more second pairs of the RF signalsengendered respectively by the one or more first pairs of the RF beams;

a transmitter, which is arranged to transmit an ablating signal towardan ablation region in the target organ via the one or more antennas; and

a processor, which is arranged to extract local tissue parameters atmultiple points in the target organ by jointly processing the RF signalsin each of the second pairs, and to cause the ablating signal to befocused on the ablation region in the target organ based on theextracted local tissue parameters.

The present invention will be more fully understood from the followingdetailed description of the embodiments thereof, taken together with thedrawings in which:

BRIEF DESCRIPTION OF THE DRAWINGS

FIGS. 1A and 1B are schematic, pictorial illustrations of radiofrequency medical imaging and therapy (RFIT) systems, in accordance withembodiments of the present invention;

FIG. 2 is a block diagram that schematically illustrates a RFIT system,in accordance with an embodiment of the present invention;

FIG. 3 is a block diagram that schematically illustrates a transmitterarray, in accordance with an embodiment of the present invention;

FIG. 4 is a block diagram that schematically illustrates a switchingarray, in accordance with an embodiment of the present invention;

FIG. 5 is a block diagram that schematically illustrates a digitalreceiver and exciter unit, in accordance with an embodiment of thepresent invention;

FIGS. 6A and 6B are diagrams that schematically illustrate a pulsegeneration circuit, in accordance with an embodiment of the presentinvention;

FIG. 7 is a diagram showing transmitted pulse sequences, in accordancewith an embodiment of the present invention;

FIGS. 8A and 8B are graphs that schematically illustrate reflectedsignal intensities measured by opposite beams, in accordance with anembodiment of the present invention;

FIG. 9 is a flow chart that schematically illustrates a method forextracting tissue properties from reflected signal intensitymeasurements, in accordance with an embodiment of the present invention;

FIG. 10 is a diagram that schematically illustrates two-dimensionalsignal reconstruction, in accordance with an embodiment of the presentinvention;

FIG. 11 is a flow chart that schematically illustrates a method forcalculating three-dimensional motion vectors, in accordance with anembodiment of the present invention; and

FIG. 12 is a block diagram that schematically illustrates a digitalreceiver and exciter unit, in accordance with another embodiment of thepresent invention.

DETAILED DESCRIPTION OF EMBODIMENTS Overview

The methods and systems described herein provide high-resolution RFimaging and therapy (RFIT) using several operational modes. These modescomprise, for example, active and passive three-dimensional (3-D)imaging modes, 3-D motion vector analysis, and RF therapy modes such asRF ablation, local heating and application of electromagnetic pressure.

In the active imaging mode, a target organ in a patient's body isirradiated with multiple RF beams generated by an antenna array. Theantenna array comprises multiple radiating elements. Subsets of elementsare selectively actuated to form multiple effective antennas, whichtransmit and receive multiple radiation beams having different azimuthsand heights. From each effective antenna (beam), the system transmits awideband signal, such as an encoded pulse sequence, having a hightemporal resolution.

The transmitted signal interacts with different tissue types in thepatient's body. Typically, RF energy is absorbed in tissue and isreflected from transition areas, i.e., interfaces between differenttissue types. Part of the transmitted signal is backscattered towardsthe effective antenna. The magnitude of the backscattered signal as afunction of time represents the tissue profile viewed from theparticular angle of the beam. For example, homogeneous tissue appears asa substantially uniform, gradually decaying magnitude. Tissue interfacesappear as temporal peaks due to the associated reflection.

The RFIT system receives and analyzes the backscattered signals in orderto image the target organ. The system may use multiple analog-to-digital(A/D) converters that sample the signal at incremental time offsets inorder to enhance the range resolution of the acquired samples. Thesystem compensates for different artifacts and measurement distortions,such as local tissue attenuation, local variations in light velocity(i.e., local time delay), and signal dispersion. Artifact compensationis performed by jointly analyzing pairs of beams that irradiate thetarget organ from opposite directions. These compensation methodsfurther improve the achievable spatial resolution of the system.

After compensating for the different artifacts, different local tissueparameters, such as attenuation, reflection, time delay, as well aslocal dielectric properties, are extracted from the measurements.

In some embodiments, measurements are performed using multiple beampairs from multiple directions. The data collected by the multiplebeams, after artifact correction, is reconstructed in both azimuth andheight, to produce a 3-D representation of the target organ. Theextracted tissue properties of the target organ are displayed in two orthree dimensions, so as to enable tissue differentiation andclassification with high spatial and temporal resolution. The systemprovides both anatomical and functional tissue information.

Unlike some known methods and systems, such as some of the known imagingmodalities cited above, the RFIT system can accurately determine thetissue type (e.g., blood, muscle, nerve, bone or fat) for each point inspace as a function of time, based on the measurement and calculation ofmultiple local tissue parameters.

In order to verify the expected performance of the disclosed methods andsystems, a series of experiments and computer simulations wereconducted, as will be described below. The methods and systems describedherein are expected to achieve a spatial resolution on the order of 1 mmper axis and a temporal resolution on the order of 100 Hz (10 ms). Thisperformance is significantly superior to the resolution of knownsystems, such as the systems and modalities cited in the backgroundsection above. As a result, the ability of a physician to diagnose andtreat various medical conditions is significantly improved.

The methods and systems described herein provide enhanced imagingperformance of dynamic organs, such as the heart or lungs, because ofthe high temporal resolution. Unlike some known imaging modalities, atemporal resolution on the order of 100 Hz enables real-time cardiacimaging. Furthermore, the high temporal resolution eliminates the needfor gated imaging, which is typically required in known imaging methodshaving slower refresh rates.

In the passive imaging mode, the system passively senses the electricalactivity of neurons, muscle cells and/or end-plates. In otheroperational modes, the system can perform high-resolution RF ablation,local heating and/or apply RF-induced electromagnetic pressure toselected tissue.

In some embodiments, a suitable contrast agent can be administered tothe patient prior to imaging, so as to enable organ-specific functionalimaging. Different system configurations may support any desired subsetof the operational modes.

Unlike some known RF imaging methods and systems, which are limited torelatively shallow penetration depths (e.g., breast imaging),embodiments of the present invention achieve high performance imaging inapplications requiring deep RF penetration, such as cardiac imaging.

System Description

FIG. 1A is schematic, pictorial illustration of a radio frequencymedical imaging and therapy (RFIT) system 20, in accordance with anembodiment of the present invention. The description that follows refersmainly to the active imaging mode, for the sake of clarity. The otheroperational modes and the system configurations supporting them aredescribed further below.

A patient 24 sits on a chair 28, which is located in the middle of acylindrical antenna array 32. The antenna array comprises multipleantenna elements 36, which are selectively combined and actuated to formmultiple effective antennas. The effective antennas transmit and receiveRF beams having different orientations to and from the patient's body,in order to image and/or apply treatment to a target organ or tissue. Inorder to perform RF ablation, local heating or generate RF-inducedpressure, however, some or all of antenna elements 36 transmit in unisontowards a certain focal point. Although the embodiments described belowmainly address cardiac imaging applications, the methods and systemsdescribed herein can be used to image and treat any other suitabletarget organ and tissue type.

In the exemplary system configuration of FIG. 1A, array 32 has adiameter of approximately 4 m and a height of approximately 80 cm. Sixhundred thirty elements 36 are distributed around the cylinderperimeter, and the array has a height of forty elements. Thus, in total,the array comprises twenty-five thousand two hundred antenna elements.In alternative embodiments, array 32 may comprise a lower or highernumber of elements and may have any other suitable shape or dimensions.In some embodiments, the system comprises one or two additionaldome-shaped antenna arrays (not shown in the figure), which arepositioned above and/or below the patient. The additional arrays furtherimprove the spatial resolution of the system, particularly whenperforming RF ablation.

Antenna elements 36 may comprise any suitable wideband radiating elementknown in the art, such as flared-notch based elements, spiral or helicalelements and horn-based elements. In some embodiments, array 32 maycomprise elements transmitting in different polarizations, so as toenable the system to perform polarization-dependent parametermeasurements. In some embodiments, elements 36 are active elements,which comprise power amplifiers for transmission, as well as low noiseamplifiers and multiple A/D converters for reception.

Chair 28 typically comprises materials that cause little distortion tothe RF radiation, i.e., materials having low reflectance and absorption.The chair may comprise, for example, polystyrene foam, wood, artificialleather and cloth. Other objects, such as various medical and surgicaltools and instruments, may be present in the vicinity of the patient.These objects should also comprise RF transparent materials or becovered with RF absorbing material.

Typically, the target organ to be imaged should be positionedsubstantially at the center of the cylindrical antenna array, in bothhorizontal and vertical dimensions. For this purpose, chair 28 istypically adjustable and may also recline so the patient can liehorizontally. The chair may have multiple adjustable degrees of freedom.In some embodiments, video cameras 40 are used for accuratelypositioning the patient at the center of the cylinder. Cameras 40 aremounted at different angles with respect to chair 28. Each camera ismounted so that the center of its field of view points to the center ofthe cylindrical array. The images produced by cameras 40 are displayedon one or more positioning displays 44. In order to position the patientcorrectly, chair 28 is remotely or locally adjusted until the region ofinterest (e.g., the patient torso) is seen at the center of the field ofview of all cameras.

Array 32 may be mounted on an elevated platform 48. Some elements ofsystem 20, such as signal generation and reception circuitry, should belocated near the antenna array in order to minimize RF losses. Suchsystem elements may be located underneath the elevated platform. A ramp52 enables wheelchair or gurney access to the platform. Other systemelements, such as signal processing elements, can be located in a rack56, located further away from array 32. Typically, a spacious area isleft around the patient, so as to allow easy access to the patient bystaff and equipment. For example, RF imaging and therapy can take placein parallel to other procedures, such as catheterization and imagingusing other modalities.

System 20 is controlled and operated from a control station 60,typically separated from array 32 by an RF absorbing wall 64. A window68 comprising RF absorbing material may be used for viewing the patientfrom the control station. The RF absorbing wall and window help toprotect staff from RF radiation. RF leakage into and out of system 20can also be reduced by covering external walls with radiation absorbingmaterial, such as RF absorbing tiles 72. The room housing system 20should be air-conditioned, in order to dissipate the heat produced bythe RF energy, particularly when performing RF ablation.

Control station 60 comprises one or more imaging displays 76, whichdisplay the imaged target organ and other relevant information. Thecontrol station also comprises input devices 80, such as a keyboard,mouse and/or trackball, for providing input and controlling the system.

FIG. 1B is a schematic, pictorial illustration of an RFIT system, inaccordance with an alternative embodiment of the present invention. Inthe system configuration of FIG. 1B, cylindrical antenna array 32 istilted at an angle typically in the range of 30-50°. Chair 28 ispositioned so that the area of interest, in the present example thetorso of patient 24, is located at the center of the cylinder.

The configuration of FIG. 1B is particularly suitable for applicationsin which staff and/or equipment are present in the vicinity of thepatient, but their effect on the RF radiation should be minimized bypositioning them outside the beam paths. For example, this configurationmay be used in different intra-operative imaging applications, such ascatheterization procedures.

In the present example, a physician 82 sits or stands below the elevatedarea of the array, by the patient's feet. A display 83 displays the dataacquired by the system to the physician. In some embodiments, radiationabsorbing clothing can be worn by the physician to minimize radiationexposure.

Operational Modes

System 20 supports several operational modes for imaging and/or therapy.Some system configurations may support all modes, whereas otherconfigurations may support only a single mode or a subset of the modes.

In some embodiments, the system supports an active imaging mode, inwhich the target organ is scanned with multiple beams. In the presentexample, the beams have horizontal beam widths of approximately 15° andvertical beam widths of approximately 4°, although other radiationpatterns can also be used. Using the exemplary array dimensions andgeometries shown in FIGS. 1A and 1B above, these beam widths aresuitable for irradiating the patient torso. In order to produce suchbeam widths, each effective antenna is approximately 15 cm wide and 50cm high. Assuming antenna elements 36 are spaced 2 cm apart, eacheffective antenna has approximately seven elements in the horizontaldimension and twenty-five elements in the vertical dimension.

The system measures multiple tissue parameters at multiple locations inthe target organ and displays them in three dimensions. Typically, thesystem directly measures three parameters for each location, namely thelocal RF attenuation, local reflection coefficient and local time-delaycaused by the decreased light velocity in tissue with respect to freespace. The measured parameters, as well as the collected raw data, mayalso be used for evaluating dielectric tissue properties, such as thelocal complex permittivity and local conductivity of the tissue.

Each of the parameters described above can be evaluated using the entiresystem bandwidth, or separately in multiple sub-bands. Parameters canalso be evaluated for different polarizations. Each parameter, or acombination of parameters, can be displayed in 3-D. For example, whenperforming cardiac imaging, the system can measure and display the localconductivity along the conduction pathways within the heart. As anotherexample, valve calcification in the heart can be accurately detectedbased on a 3-D measurement and display of local permittivity.

The multiple tissue parameters can be jointly analyzed in order toaccurately and reliably classify the tissue type, e.g., bone, muscle,fat or blood, at each location in the target organ. Using this analysis,a high-resolution display of the target organ, with each tissue typeclearly marked and differentiated, can be provided to the physician. Inthe active imaging mode, the entire heart can be imaged at a typicalframe rate of 100 Hz, without gating. The spatial resolution can reach 2mm, and often 1 mm or better.

In some embodiments, certain dynamic mechanical properties of the targetorgan can be evaluated by tracking frame-to-frame variations in theactive imaging mode. This sub-mode is referred to herein as tissuetracking. For example, tissue tracking can estimate the cardiac wallmotion velocity, as well as the local strain and local strain-rate.These properties are usually expressed as 3-D vectors.

Tracking inter-frame variations may involve known image processingtechniques, such as optic flow methods. Additionally or alternatively,anatomical landmarks can be identified in the images, either manually orautomatically. The variation in the coordinates of these landmarks canthen be tracked in different frames.

In some embodiments, a contrast agent can be used during active imaging.The contrast agent is used to produce irregular values of one or more ofthe measured parameters. For example, a contrast agent may comprise ahighly reflective substance, such as a ferrite-based substance. Somecontrast agents target a specific organ or tissue type, so as to allowhighly-specific functional imaging. For example, organ-specific agentscan be used for myocardial perfusion estimation, kidney performanceevaluation and liver performance evaluation.

In some embodiments, system 20 supports a 3-D motion vector analysismode, which measures the local motion vector for each location in thescanned target organ (e.g., the local blood velocity), as a function oftime. In this mode, the system analyzes multiple reception beamssimultaneously, and measures the Doppler shift with respect to eachbeam. The Doppler shift, as measured for each point in space withrespect to several reference points, is used to determine the dominantvelocity vector for each such point as a function of time. The fallDoppler spectrum as a function of time and space may also be calculatedfor a particular component of the vectors.

Using the motion vector analysis mode, it is anticipated that imaging ofthe entire human heart can be performed at a frame rate on the order of20 Hz, without gating. The spatial resolution in each axis is expectedto be on the order of 1 mm. The resolution of velocity measurements ineach beam is expected to be approximately 0.3 m/s, and the Nyquistfrequency (i.e., maximum unambiguous velocity) is expected to beapproximately 3 m/s.

In some embodiments, system 20 supports a non-invasive RF ablation mode.In this mode, the antenna array focuses RF energy to a small region inthe target organ in order to increase the local temperature by a factoron the order of 20° C. The local temperature can be measured in theactive imaging mode. By combining the two modes, the system canstabilize the temperature in the target spot. RF ablation can be usedfor removing tumors and cancerous cells, as well as for performingnon-invasive surgical operations such as internal hemorrhage reduction.

Typically, RF ablation in system 20 is performed in parallel to activeimaging. Unlike known ablation methods, which use different modalitiesfor imaging and ablation, in system 20 no registration is usually neededbetween the coordinate systems used for imaging and for ablation.

The spatial resolution of the ablation mode is expected to beapproximately 6 mm per axis, at the 3 dB points of the ablation region.Using higher frequency bands may allow improving the spatial resolutionby a factor of about two. When ablation is guided by active imaging, theframe rate is expected to be approximately 50 Hz, and the imagingspatial resolution is expected to be around 1 mm per axis.

Simultaneous operation of the RF ablation mode with the tissue trackingmode enables non-invasive ablation of moving organs, such as thedestruction of ectopic regions in the cardiac muscle and the removal oflung cancer cells. When RF ablation is combined with tissue tracking,the system adaptively adjusts the ablation region with a refresh ratecomparable to the imaging frame rate.

In some embodiments, system 20 can be used to applyelectromagnetically-induced pressure to target tissue. In some cases,applying instantaneous high-power electromagnetic pressure to a nervepathway may induce an action potential either directly, or due tothermal effects. Thus, applying electromagnetic pressure may help tocontrol the innate frequency of different regions within the cardiacconduction system, so as to achieve a stable sinus rhythm. The pacing ofother visceral organs, such as the gastro-intestinal system, may also beaffected by electromagnetic pressure.

Additionally or alternatively, system 20 can be used to apply localheating to tissue. Local heating can be used, for example, to treatstressed muscles and to speed the natural healing of bruised or inflamedareas. In the local heating mode, a target region can be defined andvisualized using the active imaging mode. The system can then locallyheat the selected region by applying low-power RF energy to the region.

In some embodiments, system 20 performs high-speed passive imaging ofelectrical activity in the target organ. Various cells in the humanbody, such as nerve and muscle cells, are known to show substantialelectrical activity. This activity may sometimes be detectednon-invasively by a sensitive receiver. In the passive imaging mode,system 20 triangulates the signals sensed by multiple reception beams,in order to determine the location of the electrical activity. Thepassive imaging mode is expected to reach a temporal resolution ofapproximately 1 μs and a spatial resolution of approximately 1 mm. Thisperformance level should enable the system to display electrical signalsas they pass through various physiological systems.

Data processing in system 20 can be performed either in real-time, i.e.,during data acquisition, or off-line, i.e., after data acquisition iscompleted. In the active and passive imaging modes and in the motionvector analysis mode, data processing may be performed either inreal-time or off-line. In the RF therapeutic modes, data processing istypically performed in real time in order to provide imaging guidance tothe therapeutic procedure.

System Components

FIG. 2 is a block diagram that schematically illustrates RFIT system 20,in accordance with an embodiment of the present invention. The systemcomprises a digital exciter and receiver unit 84, which generates thesignals used in the various imaging and therapy modes. The signalproduced by the exciter is split into multiple signals and amplified bya transmitter array 88. The amplified signals are distributed toindividual antenna elements 36 in antenna array 32 by a switching array92, so as to form the appropriate radiation beams. The signals aretransmitted towards the target organ by array 32.

The backscattered RF radiation is received by array 32. Switching array92 selects the appropriate subset of elements 36 corresponding to thecurrently-received beam. The signal corresponding to the specific beamis received by a digital receiver in unit 84. The signal produced by thereceiver is processed by a digital signal processor (DSP) unit 96. DSPunit 96 typically performs computationally-intensive calculations, suchas operations that are repeated many times. These calculations maycomprise, for example, compensation for artifacts and measurementdistortions and 3-D image reconstruction.

A data processor and man-machine interface (MMI) unit 100 manages thevarious real-time processes performed by system 20 and controls othersystem elements, such as transmitter array 88, digital receiver andexciter 84 and switching array 92. The management of real-time processesmay comprise mode selection, calculation of parameters to be used by DSPunit 96 and other system elements, as well as general system timing.

Unit 100 also interacts with the user, in order to accept user input andcommands. In some embodiments, unit 100 comprises a video/displayprocessor (not shown in the figure), which performs the transformationsthat translate the time-dependent 3-D images and calculated parameters,generated by DSP unit 96, to the desired viewing configurationspresented on displays 76 and 83. In some embodiments, the video/displayprocessor may also perform the final tissue classification, i.e.,determining the tissue type for each point in space based on the 3-Dimages. Additionally or alternatively, the video/display processor maycarry out tissue tracking.

Typically, DSP 96 and unit 100 comprise general-purpose or customer offthe shelf computers, which are programmed in software to carry out thefunctions described herein. The software may be downloaded to thecomputers in electronic form, over a network, for example, or it mayalternatively be supplied to the computers on tangible media, such asCD-ROM. DSP 96 and unit 100 may be implemented in a single computingplatform or in separate platforms. Some or all of the functions of DSPunit 96 may also be implemented in hardware.

In some embodiments, in particular when performing passive imaging, thethermal noise level of antenna array 32 should be reduced. For thispurpose, system 20 may comprise a cooling system 104, which coolsantenna array 32, switching array 92 and/or the receiver in unit 84.System 20 is powered by a power supply 108. Appropriate environmentalconditions, e.g., temperature and humidity, are maintained by anenvironmental control unit 112.

FIG. 3 is a block diagram that schematically illustrates transmitterarray 88, in accordance with an embodiment of the present invention. Thesignal produced by the exciter in unit 84 is split by a power splitter116 and amplified by multiple power amplifiers 120. The amplifiedsignals are provided to switching array 92. As will be shown below, thetransmitted power of each effective antenna is set to around 3 kW peakpower and 350 W average power. Assuming each effective antenna comprises25×7=175 elements, the peak power of each amplifier 120 is approximately17 W and the average power is 2 W. These power levels are readilyachievable using known solid-state devices.

Each amplifier 120 produces a signal that will ultimately drive aparticular element 36 in the currently-used effective antenna. Thus, thenumber of amplifiers 120 in the transmitter array should match thenumber of elements in the effective antenna. In alternative embodiments,a single amplifier can be allocated to a group of elements in theeffective antenna.

In the RF ablation mode, the number of elements assigned to a particularamplifier may depend on the desired ablation region size and on theability to control the antenna array focal point. In the passive imagingmode, the transmitter array is not used.

In some embodiments, the beam of the effective antenna is focused andshaped by applying different weights and/or different timing offsets tothe different elements. This process is commonly referred to asapodization. In these embodiments, transmitter array 88 comprises atiming and apodization module 124, which is controlled by the dataprocessor. Module 124 adjusts the timing, gain and/or phase ofamplifiers 120, in accordance with the apodization scheme used. In someembodiments, encoding of the transmitted pulse sequence, a function thatis described in detail further below, is also carried out by module 124.

FIG. 4 is a block diagram that schematically illustrates switching array92, in accordance with an embodiment of the present invention. Switchingarray 92 is connected to antenna array 32. An exemplary effectiveantenna 128 is shown in the figure. On transmission, switching array 92accepts the signals from transmitter array 88 and routes them to theappropriate elements 36 in antenna array 32, in accordance with thecurrently-used effective antenna. On reception, switching array 92routes the signals received by the elements of the currently-usedeffective antenna to the receiver in unit 84.

Switching array 92 comprises a switch matrix 132, which selects theappropriate subset of elements 36. The switch matrix is controlled bythe data processor of unit 100. The switching array further comprisesmultiple duplexers 136. On transmission, the duplexers isolate thetransmitted signals from the receiver. In some embodiments, the geometryand pulse repetition frequency (PRF) of the system are configured sothat transmission and reception do not occur simultaneously. In theseembodiments, T/R switches or circulators may be used instead ofduplexers 136.

The operation and/or configuration of the switching array may vary fordifferent operational modes of system 20. For example, in the activeimaging mode, a single receiver channel may be assigned to each elementof the currently-used effective antenna, to a group of elements, or tothe entire effective antenna. In the motion vector analysis mode, asingle transmit beam and several (e.g., 10) receive beams are used. Thereceive beams are often narrower than in the active imaging mode,therefore the effective antennas comprise a higher number of elements36. During RF ablation, most or all elements participate in thetransmission, and no reception is performed. In the passive imagingmode, no transmission is carried out, and several receive beams areused, either alternately or simultaneously.

In order to achieve the desired frame refresh rate, system 20 performshigh-speed beam switching, typically on the order of 9 MHz (i.e.,scanning of 9,000,000 beams per second). The switching elements inswitch matrix 132 should support this high switching speed. For example,PIN-diode switches, such as the S9H-79-3 device produced by GTMicrowave, Inc. (Randolph, N.J.), can be used for this purpose. Thesedevices have a 30 ns switching time. Further details regarding thesePIN-diode switches are available at www.gtmicrowave.com.

FIG. 5 is a block diagram that schematically illustrates digitalreceiver and exciter unit 84, in accordance with an embodiment of thepresent invention. Unit 84 comprises an exciter 140, which produces thepulsed signal waveforms used for transmission and provides the signalsto transmitter array 88. As shown in FIG. 2 above, on reception, unit 84accepts the received signals from elements 36 of the currently-usedeffective antenna via switching array 92. The signals are combined usinga power combiner 142.

A correlator module 144 correlates the received signal with the expectedsignal, i.e., the signal waveform produced by the exciter. The pulseshape of the transmitted pulses is typically selected so that itstemporal point spread function (PSF) has low sidelobes. In someembodiments, module 144 performs additional functions such as matchedfiltering, down-conversion to a suitable baseband or intermediatefrequency (IF) and/or pulse de-compression. Module 144 may also applytime-dependent gain control (TGC) to the signal.

In alternative embodiments, pulse de-compression can be performed usingsuitable software or digital hardware after the signal is digitized.Signal down-conversion may also be performed on the digitized signal. Insuch configurations, the signal should be digitized at a sampling ratecorresponding to the highest radio frequency used (e.g., 18 GHz).

As will be described in greater detail below, the range resolution ofthe system is achieved by sampling (digitizing) the signal usingmultiple analog-to-digital (A/D) converters that sample the signal atincremental time offsets. For this purpose, the signal produced bymodule 144 is split by a 1:15 power splitter 146. The fifteen outputs ofthe power splitter are delayed by fifteen delay lines 148.

In the exemplary system configuration described herein, the signalproduced by module 144 enables a raw spatial resolution of 1.5 cm. Delaylines 148 divide this range into 15 intervals. In other words, the delaydifference between successive delay lines is equivalent to a 1 mm range.The outputs of the delay lines are sampled by fifteen synchronized A/Dconverters 150. The sampled signals are provided to DSP unit 96.

The exemplary configuration of FIG. 5 is typically suitable for theactive imaging mode. In the passive imaging mode, the system does nottransmit. In this mode, module 144 correlates the received signal with asynthetically-produced signal that approximates the signal waveform thatis expected to be produced by the target tissue. In the RF ablationmode, exciter 140 produces the ablating signal waveform, and thereceiver is used only when performing imaging. In the motion vectoranalysis mode, the target organ is imaged simultaneously by severalbeams. An exemplary receiver configuration suitable for this mode isdescribed in FIG. 12 further below.

In alternative embodiments, some of antenna elements 36 can be definedas transmit-only elements, and other elements may be defined asreceive-only elements. Such configurations reduce the number ofduplexers and cables, and simplify the system calibration. Hybridconfigurations in which some elements are transmit-only, some arereceive-only and some perform both transmission and reception, are alsofeasible.

In some embodiments, the system configuration can be simplified byrelaxing some of the system requirements. For example, the system can bedefined to support only a single operational mode or a small subset ofmodes. Defining the system for a smaller penetration depth and/or slowerrefresh rate can also simplify the system. For example, when using aslower refresh rate, the system can use transmitted signals havingsimpler waveforms, such as waveforms based on stepped frequency, linearfrequency modulation, complementary coding and other phase codingmethods. The system can use longer pulse sequences in conjunction withthese waveforms.

In alternative embodiments, the system can perform mechanical antennascanning instead of electronic scanning. For example, a single antennacan be steered mechanically around the patient in one or more axes, soas to produce the multiple beams needed for imaging. As another example,a pair of antennas can be positioned on opposite sides of the patientand steered mechanically around the patient, forming pairs of beams onopposite sides of the patient. Alternatively, two or more antenna pairscan be used. Further alternatively, the antenna or antennas can bestationary, and the patient can be moved and/or rotated with respect tothe antennas.

Additionally or alternatively, the distance between the antenna and thepatient can also be changed incrementally by moving either the antennaor the patient. Using multiple distances between the antenna and patientis equivalent to using multiple A/D converters having incremental timeoffsets. Spherical scanning, or any other suitable scanning geometry,can be used instead of cylindrical scanning, when using eithermechanical or electrical scanning.

In system configurations based on mechanical scanning, the number ofantennas and amplifiers is significantly reduced, and the switchingarray can be significantly simplified or altogether eliminated. Suchconfigurations may be particularly suitable for applications that do notinvolve hospitalization, such as in dentistry, plastic surgery,ophthalmology and orthopedic applications. Configurations in which theantennas are stationary and the subject is moved may be useful inexperimental and non-medical applications, such as in animalexperiments.

Spatial Resolution

The spatial resolution expected to be achieved by system 20 in theactive imaging mode is on the order of 1 mm per axis. In particular,when processing the backscattered signal of a certain beam, the systemhas a range resolution of 1 mm. This resolution level is achieved by acombination of (1) using a wideband transmitted waveform having a rangeresolution of ˜1.5 cm, and (2) dividing the 1.5 cm resolution into 1 mmeffective range gates by processing the received signal using fifteenA/D converters having incremental time offsets.

The transmitted signal produced by exciter 140 comprises a sequence ofnarrow pulses (narrow in time and wide spectrally). Each pulse has abandwidth of approximately 10 GHz, i.e., a pulse width of approximately0.1 ns. Typically, the spectral content of the pulse covers the range of8-18 GHz. In alternative embodiments, higher bandwidths, such as 6-25GHz, may also be feasible.

In order to improve signal-to-noise (SNR) ratio, exciter 140 produces asequence of sixty-four successive pulses. The sequence is phase-encoded,i.e., each pulse in the sequence is given a certain phase shift. In thepresent example, bi-phase encoding is used, in which the phase shiftsare either 0° or 180°. Alternatively, any other suitable encoding schemecan be used. On reception, the received pulse sequence is correlatedwith a reference sequence, so as to achieve the desired pulsecompression gain.

Narrow, wideband pulses can be produced, for example, by chopping anarrowband signal whose frequency approximately matches the desiredtransmit carrier frequency. Alternatively, a baseband signal can bechopped and then up-converted to the desired transmit frequency. Highspeed chopping can be performed, for example, using step recovery diodes(SRD). For example, Aeroflex/Metelix Inc. (Sunnyvale, Calif.) offers asilicon SRD device denoted MMDB30-B11, which can be used for thispurpose. These diodes are capable of generating 30 ps pulses. Furtherdetails are available at www.aeroflex-metelics.com.

Since the recovery time of SRD devices is relatively long, a sequence ofshort pulses can be generated by multiple SRD devices in parallel, whichare actuated sequentially. Each SRD generates a single pulse in thesequence at the appropriate timing.

FIGS. 6A and 6B are diagrams that schematically illustrate a pulsegeneration circuit 152, which may be used by exciter 140 to generate thetransmitted pulse sequences, in accordance with an embodiment of thepresent invention. Circuit 152 of FIG. 6A comprises a sequence ofSRD-based switches 154. Circuit 152 shows only eight switches for thesake of simplicity, however similar circuits having different numbers ofswitches can be used to generate pulse sequences having any desirednumber of pulses. Each switch 154 has an input 156, which accepts aphase-encoded RF signal. In the present example, bi-phase encoding isused. The input to a switch whose pulse is encoded with a 0° phase shiftis marked “+”, and the input to a switch whose pulse is encoded with a180° phase shift is marked “−”. The switches are actuated sequentiallyby a timing logic circuit 158. The SRD outputs are combined using acombiner 160 to produce the desired pulse sequence.

FIG. 6B shows an exemplary circuit for generating the phase-encoded RFsignal used as input to the different SRD switches, in accordance withan embodiment of the present invention. An oscillator 162 generates acontinuous sinusoidal signal at the desired transmit frequency. Theoutput of oscillator 162 is split by a power splitter 164. One output ofthe splitter is provided to the SRD stages encoded with a 0° phase shift(the stages marked with “+” in FIG. 6A). The other output of thesplitter is phase-inverted using a 180° phase shifter 166, and providedto the SRD stages encoded with a 180° phase shift (the stages markedwith “−”).

FIG. 7 is a diagram showing transmitted pulse sequences generated byexciter 140, in accordance with an embodiment of the present invention.In the present example, exciter 140 transmits sequences of 64phase-encoded pulses. Each pulse is 0.1 ns wide, so that the overalltransmitted signal has a length of 6.4 ns. The exciter transmits pulsesequences at a pulse repetition interval (PRI) of 55 ns. This PRIcorresponds to a two-way range of over 8 m in free space, significantlymore than the radius of antenna array 32, thus avoiding range ambiguity.In some embodiments, two or more pulse sequences are transmitted perbeam. In these embodiments, for each beam position, the data of eachrange gate may be integrated over the different pulses in order toenhance SNR. The 55 ns PRI is selected in order to achieve a temporalresolution of 100 Hz, as will be described in greater detail in FIG. 10below.

As noted above, the reflected signal is sampled by fifteen parallel A/Dconverters having incremental time offsets equivalent to 1 mm. Thefifteen A/D converters effectively divide the 1.5 cm range gatesachieved by the wideband pulses into effective range gates of 1 mm. Insome embodiments, DSP unit 96 solves a set of linear equations based onthe outputs of the fifteen A/D converters, and evaluates the receivedsignal with a resolution of 1 mm. The DSP unit may apply knowndeconvolution methods for this purpose. This process is typicallyperformed separately for each pulse in the pulse sequence. In someembodiments, after processing each pulse, DSP unit 96 integrates thedata for each range gate over the sixty-four pulses in the sequence inorder to improve the measurement SNR.

Module 144 in the receiver performs time-dependent gain control (TGC)prior to digitization of the signal, in order to provide sufficientdynamic range at the A/D converters. The TGC process typically uses afixed or pre-calibrated function that specifies the attenuation as afunction of range. The function may be evaluated by occasionallytransmitting a narrow calibration beam and measuring the attenuation asa function of range. Typically, the TGC process provides coarse gaincontrol, and an additional fine gain control process is performeddigitally, after the signal is sampled by the A/D converters.

The description above refers to a cardiac imaging application, which ishighly-dynamic and imposes harsh refresh rate requirements. For organsother than the heart, a 100 Hz refresh rate may not be necessary. Inthese cases, longer pulse sequences, which may comprise up to 1000pulses or more, may be used. Using long pulse sequences significantlyincreases the achievable penetration depth.

Parameter Extraction and Measurement Artifact Compensation

The reflected signal measurements performed by system 20 are oftendistorted as a result of the different physical properties of the imagedtissue. Different tissue types differ from one another, and from freespace, by their light velocity, local attenuation and signal dispersionproperties. In some embodiments, system 20 compensates for theseartifacts in order to achieve high spatial resolution.

As noted above, system 20 scans the target organ using multiple beamsfrom multiple directions. In some embodiments, the system compensatesfor tissue artifacts by jointly-analyzing the reflections measured bypairs of beams that image the target organ from opposite directions.

FIGS. 8A and 8B are graphs that schematically illustrate reflectedsignal intensities measured by opposite beams, in accordance with anembodiment of the present invention. In FIG. 8A, a curve 220 shows thereflection intensity as a function of range, as measured by a particularbeam. Curve 220 shows three characteristic peaks 224A, 224B and 224C,which are typically produced by tissue transitions (i.e., interfacesbetween different tissue types). Peak 224A is a strong peak having theshortest range to the effective antenna. A peak of this kind is usuallyproduced by the body “skin effect,” i.e., the transition from free spaceto tissue when penetrating the skin. Peaks 224B and 224C relate to othertransitions from one tissue type to another.

In FIG. 8B, a curve 228 shows the reflection intensity as a function ofrange, as measured by another beam, which is located at a 180° azimuthoffset with respect to the beam of FIG. 8A. In other words, the beam ofFIG. 8B images a similar depth cross section of the target organ, butfrom the opposite direction. Note that the range (horizontal) axis ofFIG. 8B is reversed, so as to enable curves 220 and 228 to be compared.Curve 228 also shows three peaks denoted 224D, 224E and 224F. Peaks 224D. . . F correspond to the same tissue transitions as peaks 224A . . . c,respectively, as seen from opposite directions.

In FIGS. 8A and 8B the range coordinates of the peaks in the twoopposite beams are shown to coincide with one another. In many practicalcases, however, the measured coordinates of the peaks generally differfrom one another in the opposite direction measurements because of thedifferences in local light velocity (local time delay) across thetissue. These differences are used to estimate and compensate for thelocal tissue time delay.

FIG. 9 is a flow chart that schematically illustrates a method forextracting tissue properties from reflected signal intensitymeasurements of opposite beams, in accordance with an embodiment of thepresent invention. The method of FIG. 9 refers to a particular depthcross section of the target organ, as imaged by a pair of beams at a180° offset. The process of integrating the data measured by multiplebeam pairs at different azimuth angles and heights into a 3-Drepresentation of the target organ is described further below.

The method begins with system 20 acquiring raw reflection measurementsusing two beams at opposite directions, at a data acquisition step 240.The intensity measurements as a function of range acquired by aparticular beam are referred to hereinbelow as a measurement set. Thesystem enhances the range resolution of the two measurement sets byusing multiple A/D converters, as described above, at a range resolutionenhancement step 244. After reversing the range dimension of one of themeasurement sets, the reflection measurements of the two beams resemblethe exemplary reflection measurements of FIGS. 8A and 8B above.

In many cases, the skin-effect reflection, i.e., the reflection peakhaving the shortest range to the effective antenna, is substantiallystronger than other, deeper reflections. Temporal side-lobes generatedby this peak may distort or mask other reflection peaks. The side-lobesare a result of the point spread function (PSF) of system 20. In someembodiments, the distortion caused by the skin-effect peak can bereduced by detecting the true position of the skin-effect peak, andsubtracting from the measured signal a replica of the system PSF,centered at the measured skin-effect peak location. This process istypically performed for each beam position. The skin surface may beadded after carrying out the different artifact correction procedures.

System 20 performs time delay estimation and compensation, at a timedelay correction step 248. The system identifies in the two measurementsets pairs of characteristic peaks that correspond to the same tissuetransition (interface), as viewed from the two opposite directions (suchas, for example, the pair 224A and 224D in FIGS. 8A and 8B above). Asnoted above, the measured peak locations are offset from one another inthe two measurement sets due to local time delay differences. For eachpeak, the system calculates an estimated peak location based on the peaklocations in the two measurement sets.

The system then applies piecewise translation and stretchingtransformations to the horizontal axes of the two measurement sets,which moves the peaks to the estimated locations. Thus, each peak isassociated with two time delay offsets, i.e., the time offsets betweenthe original peak location in the two measurement sets and the estimatedlocation. These time offsets are proportional to the cumulative localtissue time delay from the effective antenna and up to the location ofthe peak. In between peaks, the tissue is assumed to be relativelyhomogeneous, therefore the time delay offset is distributed uniformlybetween successive peaks.

For each pair of characteristic peaks, the estimated peak location canbe determined as follows: The ranges, measured from the direction of oneof the two effective antennas, corresponding to the i′th characteristicpeak in the two measurement sets, are denoted r_(i) ¹ and r_(i) ²,respectively. The estimated location of the i′th peak is denoted r_(i).The time delay, with respect to the timing corresponding to the samerange in free space, which is caused by the tissue between the (i−1)′thpeak and the i′th peak, is denoted d_(i), when expressed in units ofrange. Using this notation, it can be shown that:

$r_{i}^{1} = {r_{i} + {\sum\limits_{1}^{i}d_{i}}}$$r_{i}^{2} = {r_{i} - {\sum\limits_{i + 1}^{P}d_{i}}}$wherein P denotes the number of characteristic pairs of peaks. This setof linear equations can be solved to extract both r_(i) and d_(i). Thesecalculations assume there is no time delay outside the subject body.

After completing step 248, the system has an estimate of the localtissue time delay for each point (effective range gate) along the crosssection of the target organ.

System 20 estimates the local tissue attenuation, at an attenuationestimation step 252. First, the system calculates the total side-to-sidebody attenuation along the examined depth cross section. The systemidentifies in the two measurement sets a pair of peaks that correspondto a particular skin transition (e.g., the pair 224A/224D or the pair224C/224F in FIGS. 8A and 8B above). The ratio between the reflectionintensities of the two peaks is used as an estimate of the total bodyattenuation.

Using the known total attenuation, the cumulative attenuation from theskin and up to a particular peak can be calculated. The cumulativeattenuation is calculated by considering the following four equalities:(1) the total body attenuation is identical in the two measurement sets;(2) the measured peak intensity in the first measurement set (i.e.,measured from the direction of the first effective antenna) is equal tothe true reflection at the peak location, plus the cumulative tissueattenuation from the skin and up to the peak from this direction; (3)the measured peak intensity in the second measurement set (i.e.,measured from the direction of the second effective antenna) is equal tothe true reflection at the peak location, plus the cumulative tissueattenuation from the skin and up to the peak from the oppositedirection, and (4) the sum of the two cumulative attenuations (from theskin and up to the peak) from the two opposite directions is equal tothe known total body attenuation.

The system solves the linear equations derived from these fourequalities, and extracts the true reflection coefficient of each peak,as well as the cumulative attenuation between every two successivepeaks. The local tissue attenuation and reflection are assumed to beuniform between peaks. Therefore, the cumulative attenuation isdistributed uniformly between peaks. The system adjusts the intensity ofthe measurement sets, so as to compensate for the local attenuation.After completing step 252, the system has estimates of the local tissueattenuation and reflection coefficients for each point (effective rangegate) along the cross section of the target organ.

Note that when measuring and manipulating peak intensities, any gaincontrol applied by the receiver should be taken into account. Forexample, when the receiver employs time-dependent gain control (TGC),the TGC value of each range gate should be taken into consideration, inorder to measure the peak intensities correctly.

In some embodiments, the accuracy of the reflection coefficientcalculation can be improved by averaging the two attenuation-correctedmeasurement sets. Additionally or alternatively, the estimation accuracycan be improved by performing several iterations of time delay andattenuation correction. In some embodiments, a background signal perrange gate, typically pre-calibrated without the presence of tissue, canbe subtracted from the corrected measurement of each range gate. Anexemplary pre-calibration process is described further below.

In some embodiments, the estimation process of steps 240-252 above canbe performed separately in several spectral sub-bands in order to reducethe effects of signal distortion. For example, the 10 GHz signalbandwidth can be divided into 5-10 sub-bands of 1-2 GHz bandwidth, andthe local time delay, reflection and attenuation calculated separatelyfor each sub-band. In order to process each sub-band, the reflectedsignal should be filtered, so as to retain only the spectral content inthe currently-processed sub-band. Such filtering can be performed eitherbefore or after the A/D converters, using either analog or digitalfiltering, respectively. If desired, the corrected measurement sets canbe summed over all sub-bands to produce equivalent wideband measurementsets.

The system now uses the three estimated local tissue parameters (localtime delay, reflection and attenuation) to evaluate the local dielectricproperties along the examined depth cross section, at a dielectricevaluation step 256. Several models are known in the art for calculatingdielectric properties based on such physical parameters. For example, asingle dielectric layer model is described by Yariv in “OpticalElectronics,” CBS College Publishing, 1985, Chapter 4, pages 87-95,which is incorporated herein by reference.

The single layer model cited above is particularly suitable forcontinuous transmission. In some embodiments, a more complex model thatconsiders multiple dielectric layers and pulsed transmission can beused. The model analyzes a structure of multiple adjacent layers havingdifferent material composition, with the first layer representing freespace. For example, consider a structure of three layers and threetransitions (air to first layer, first to second layer, second to thirdlayer). The three layers have dielectric constants denoted ∈₂, ∈₃, and∈₄. ∈₁=∈₀ represents the dielectric constant of air. At the k′th layertransition, the incident wave is denoted I_(k) ^(L), and the waveprogressing to the next layer is denoted I_(k) ^(R). The incident wavecomponents can be written asI₁ ^(L)=1I ₁ ^(R)=|(1+r ₁)|I ₂ ^(L)=|(1+r ₁)exp(jk ₂Δ₂)|I ₂ ^(R)=|(1+r ₁)exp(jk ₂Δ₂)(1+r ₂)|I ₃ ^(L)=|(1+r ₁)exp(jk ₂Δ₂)(1+r ₂)exp(jk ₃Δ₃)|I ₃ ^(R)=|(1+r ₁)exp(jk ₂Δ₂)(1+r ₂)exp(jk ₃Δ₃)(1+r ₃)|wherein r_(i) denotes the reflection coefficients of the i′th layertransitions, Δ_(i) denotes the thickness of the i′th layer, and k_(i)denotes the wave number in the i′th layer.

At each layer transition there exist reflected (outgoing) wavecomponents from both the present layer transition and from deeper layertransitions. In general, a reflected component originating from the i′thlayer transition and having traversed (j−1) layers after being reflectedis denoted O_(i) ^(j). The outgoing components are given by

O₁¹ = r₁ O₂¹ = (1 + r₁)exp (j k₂Δ₂)r₂O₂² = (1 + r₁)exp (j k₂Δ₂)r₂exp (j k₂Δ₂)(1 + r₁)O₃¹ = (1 + r₁)exp (j k₂Δ₂)(1 + r₂)exp (j k₃Δ₃)r₃$O_{3}^{2} = {\begin{matrix}{\left( {1 + r_{1}} \right){\exp\left( {j\; k_{2}\Delta_{2}} \right)}\left( {1 + r_{2}} \right){{\exp\left( {j\; k_{3}\Delta_{3}} \right)} \cdot}} \\{r_{3}{\exp\left( {j\; k_{3}\Delta_{3}} \right)}\left( {1 + r_{2}} \right)}\end{matrix}}$ $O_{3}^{3} = {\begin{matrix}{\left( {1 + r_{1}} \right){\exp\left( {j\; k_{2}\Delta_{2}} \right)}\left( {1 + r_{2}} \right){{\exp\left( {j\; k_{3}\Delta_{3}} \right)} \cdot}} \\{r_{3}{\exp\left( {j\; k_{3}\Delta_{3}} \right)}\left( {1 + r_{2}} \right){\exp\left( {{jk}_{2}\Delta_{2}} \right)}\left( {1 + r_{1}} \right)}\end{matrix}}$

The total outgoing signal from the first layer transition is equal to O₁¹+O₂ ²+O₃ ³. O_(n) ^(n) can be written as

$O_{n}^{n} = {A\;{{\exp\left( {{{- j}\;{wt}_{n}} + \phi} \right)} \cdot r_{n} \cdot \left\lbrack {\prod\limits_{m = 2}^{n}\;{\exp\left( {2\; j\; k_{m}\Delta_{m}} \right)}} \right\rbrack \cdot \left\lbrack {\prod\limits_{m = 1}^{n - 1}\;\left( {1 + r_{m}} \right)^{2}} \right\rbrack}}$wherein A denotes the incident signal magnitude, φ denotes the phase atthe first layer transition, ω denotes the signal angular velocity, andt_(n) denotes the time index at which the signal was received,corresponding to the two-way path of the beam. Using this model, thedielectric coefficients ∈_(i) can be extracted from the measuredresults.

The models described above assume a planar incident wave front. Inalternative embodiments, the model may be adapted to take into accountthe spherical decay of spherical wave fronts, such as by dividing thesignal by R². In some cases, the incident wave is not perpendicular tothe plane of layer transition, an effect which may also cause theresults to be polarization-dependent.

When applying any suitable layer model to the measurements of system 20,the layer thicknesses correspond to distances between successive peaks,after attenuation and time delay correction. Additionally, the localattenuation and reflection coefficients are also known, as describedabove. The complex permittivity of each layer can be calculated(assuming perpendicular incidence) using the well-known equations

${r_{n}} = {\frac{\sqrt{ɛ_{n + 1}} - \sqrt{ɛ_{n}}}{\sqrt{ɛ_{n + 1}} + \sqrt{ɛ_{n}}}}$${k_{n}} = {{k_{0}\sqrt{\frac{ɛ_{n}}{ɛ_{0}}}}}$wherein r_(n) and k_(n) respectively denote the reflection andattenuation of the n′th layer. The imaginary component of thepermittivity, denoted ∈″, and the conductivity, denoted σ, roughly obeythe relationship ∈″=σ/ω. Therefore, the conductivity σ can be derivedfrom the estimated permittivity.

An alternative method for estimating the complex permittivity may bebased on the signal within each layer separately. Assuming no losses,the light velocity v in a particular layer is given by:

$v = {\frac{\omega}{k} = {\frac{c}{n} = \frac{1}{\sqrt{\mu\; ɛ}}}}$wherein n denotes the index of refraction, and ω is the signal angularvelocity. In the presence of losses, k is complex. Defining k as

${k \equiv {\beta - {j\frac{\alpha}{2}}}},$α and β are given by the equations

${\beta^{2} - \frac{\alpha^{2}}{4}} = {\frac{\omega^{2}}{c^{2}}{{Re}\left( ɛ_{r} \right)}}$${\beta\;\alpha} = {\frac{\omega^{2}}{c^{2}}{{Im}\left( ɛ_{r} \right)}}$

This model is described in detail by Jackson in “ClassicalElectrodynamics,” John Wiley & Sons Inc., New York, 1999, pages 295-316,which is incorporated herein by reference.

α and β may be estimated by measuring the complex signal (beforeattenuation correction and time-delay compensation) immediately before areflection peak, and comparing it to the signal at a previous reflectionpeak. Using the estimated α and β values, both the real and imaginarycomponents of the permittivity can be calculated. Note that this modelis particularly suitable for narrowband signals, such as when the signalis confined to a particular sub-band of system 20.

Additionally or alternatively, any other suitable model can be used toestimate the dielectric tissue properties based on the measured timedelay, reflection and attenuation values.

In summary, the method of FIG. 9 provides estimates of local tissue timedelay, reflection and attenuation coefficients, as well as localdielectric properties, based on backscattering intensity measurementsmade by two beams that irradiate the target organ from oppositedirections. Using this method, the reflected signal intensitymeasurements acquired by the different beams are corrected, to accountfor local time delay and attenuation.

Most of the parameters measured by system 20 do not typically depend onthe geometry of the measurement. For example, Gabriel et al., in “TheDielectric Properties of Biological Tissues: I. Literature Survey,”Physics in Medicine and Biology, volume 41, 1996, pages 2231-2249, whichis incorporated herein by reference, refer to the dielectric propertiesof tissue as scalars and not as tensors. Reflection coefficientmeasurements, on the other hand, often do depend on the geometry of themeasurement, e.g., on the angle of incidence with respect to the tissueinterface that produces the reflection. In system 20, however, everypoint in space is measured from multiple directions. As a result, thegeometry-dependent effects are inherently averaged and minimized.

The possible spatial dependency of the reflection coefficient, eitherdue to angular dependency of the dielectric properties or due to theeffect of the tissue interface geometry, can be used to extractadditional clinical information. For example, 3-D reconstructed imagesmay be produced using multiple beam configurations, with the beamconfiguration changing cyclically from one frame to another. Thedifference between successive frames may be indicative of the spatialdependency. The beam configurations may differ from one another inseveral aspects, such as the elevation of the beam with respect to theplane of the cylinder.

Scanning and Reconstruction

System 20 integrates the extracted tissue parameters measured by thedifferent beams, after artifact compensation, into 3-D representationsof these parameters. In principle, a scanned volume containing thetarget organ is divided into horizontal slices, at increments on theorder of 1 mm. Each horizontal slice is scanned from multipledirections. Artifacts are compensated for, and parameters are extractedusing pairs of opposite beams, as described above. The system thenperforms two-dimensional (2-D) reconstruction of the reflected signalintensity across the slice, using the data collected by the multiplebeams.

FIG. 10 is a diagram that schematically illustrates two-dimensionalsignal reconstruction in a particular horizontal slice, in accordancewith an embodiment of the present invention. Although the descriptionthat follows refers to the 2-D reconstruction of reflected signalintensity values, the method is similarly used for reconstructing theother extracted tissue parameters across the slice. In other words, themethod can be used to reconstruct the attenuation, reflection and/ortime delay values across the 2-D slice.

The time delay map may be directly translated into a map of local lightvelocities. The values of the local dielectric properties, as a functionof time and space, may be derived from the above three reconstructedimages (i.e., the attenuation map, the reflection map and the time delaymap).

The diagram shows a top view of a particular slice. The area of theslice is divided by a grid 200 into multiple grid cells 204. Two beams206 and 207 view the slice from two different directions. The twoangular sectors shown in the figure indicate the approximate horizontalbeam-widths of the beams.

As described above, for each beam, system 20 measures the reflectedsignal intensity in 1 mm effective range gates, which are then correctedto account for local time delay and attenuation. In beam 206, aparticular effective range gate 208 corresponds to an arc having a widthof Δr=1 mm. A similar effective range gate 212 is shown for beam 207. Ascan be seen in the figure, each effective range gate covers multiplegrid cells 204.

In order to perform 2-D reconstruction of the reflected signalintensity, DSP unit 96 accumulates the intensities contributed to eachgrid cell 204 by the different beams and effective range gates. For eacheffective range gate of each beam, the system previously measured aparticular intensity value. This intensity value is divided uniformlyamong the grid cells covered by the effective range gate in question.The process is repeated for all grid cells, beams and effective rangegates. At the end of the process, each grid cell has an accumulatedvalue, integrated over the different effective range gates of thedifferent beams scanning the particular slice.

The 2-D reconstruction process described above refers to a particularhorizontal slice. The process is repeated for all slices, to produce a3-D grid. Although the vertical increment between neighboring slices ison the order of 1 mm, the thickness of each slice is derived from theelevation beam-width of the effective antennas used, typically on theorder of 15 cm. Thus, there is significant overlap between the slices.

System 20 translates the measurements performed in the overlappingslices into a set of linear equations. Solving the equations achieves aspatial resolution on the order of the increment size (in the presentexample 1 mm) in the vertical dimension. The equations may be solved,for example, using deconvolution methods. Deconvolution methods areoften noise-sensitive. Therefore, a sufficiently high SNR should beobtained.

Accurately solving the equations often involves knowledge of thesolution at the boundaries of the scanned region. In some embodiments,protective RF absorbing sheets having a negligible reflectioncoefficient can be placed at the upper and lower boundaries of thesubject's relevant body area, so as to force known boundary conditions.Thus, the equations may be solved analytically.

In some embodiments, the contribution of each effective rage gate can bedivided non-uniformly among the grid cells covered by the effectiverange gate. For example, the distribution can be range-dependent.Alternatively, the distribution can take into account angle-dependentgain differences of the effective antenna. Other known methods, such asiterative back-projection and filtered back-projection, which aresometimes used in tomography systems, can also be used for improving thespatial resolution in system 20.

In the description above, effective range gates are geometricallyrepresented as concentric arcs centered at the phase center of theeffective antenna. In many practical cases this representation isinaccurate, since the light velocity in tissue is different from thelight velocity in free space, and also between different types oftissue. As a result, the actual geometrical shape of an iso-time surface(i.e., the locus of all points having a certain propagation delay fromthe effective antenna) deviates from a perfect arc. Estimating theactual shape of the iso-time surfaces produces a more accuraterepresentation of the effective range gates.

For example, the iso-time surfaces can be estimated by considering twodifferent light velocity values, a free space value and a representativelight velocity value in tissue. The shape of the iso-time surfaces canbe determined based on an estimation of the shape of the patient's outerbody surface. Points determined to be within the patient body areassigned the tissue light velocity value, and points determined to beoutside the body are assigned the free space value.

The shape of the outer body surface can be estimated, for example, usingthe following process:

-   -   For each beam, measure the minimal distance at which a        non-negligible backscattering intensity is received (i.e.,        measure the minimal distance from the effective antenna phase        center to the body surface).    -   The volume confined within the cylindrical antenna array is        represented by a 3-D array, whose values are initially set to        zero. For each beam, all the elements whose distance from the        effective antenna phase center matches the measured distance to        the surface, and that are within the main-lobe of the respective        beam, are set to “1”. The resulting set of coordinates for each        beam typically resembles a dome-shaped surface.    -   Assume that the point at the 3-D center of the cylinder is        located within the scanned tissue. Let us now examine the values        of the 3-D array along radial lines originating from the central        of the cylinder and pointing to different 3-D angles        homogenously spanning the 4π sphere. Along each such radial, the        outer surface of the patient body is defined by the 3-D array        element whose value is “1” and its distance from the center of        the cylinder is minimal.

Additionally or alternatively, the iso-time surfaces can be adjustedaccording to the local time delay at each point within the patient body.This information is typically available after the reconstruction process(based on outer body surface estimation) is completed. Thereconstruction process inherently produces a 3-D map of local timedelays in the scanned volume, which can be used to recalculate theiso-time surfaces for each beam. The time delay values can then be usedto refine the reconstructed image. Such a procedure may be repeatediteratively, for example until the incremental variations in eachiteration become negligible.

In the present example, system 20 scans a volume having a height of 25cm, a size typical of cardiac imaging. The volume is divided into 250horizontal slices at 1 mm increments. Each slice is scanned by 360beams, which are distributed at 1° increments around the perimeter ofthe cylindrical antenna array. In each beam, two 64-pulse sequences aretransmitted. Thus, a total of 250×360×2=180,000 pulse sequences aretransmitted in order to scan the volume once. In order to achieve atemporal resolution of 100 Hz, system 20 transmits 18,000,000 pulsesequences per second, yielding the PRI of 1/18,000,000=55 ns shown inFIG. 7 above.

In the present example, the height dimension of the imaged volume isdivided into 250 slices using an array having forty antenna elementsalong the height dimension. When imaging a particular slice, the beamsare typically shaped so that their phase centers fall in the plane ofthe slice. As a result, some of the beams may not be perfectly parallelwith the base of the cylinder, i.e., some beams may be slightly tiltedin elevation. The tilt is typically on the order of several tenths of adegree, usually no greater than one degree.

In alternative embodiments, other scanning and reconstruction processescan be used. For example, instead of dividing the scanned volume intoslices, the vertical dimension can be reconstructed similarly to thehorizontal reconstruction method of FIG. 10, since each beam has acertain width in elevation. In general, the volume can be divided into athree-dimensional grid comprising multiple 3-D grid cells. The reflectedsignal intensity of each 3-D grid cell can be evaluated by scanning overbeams in azimuth as well as elevation.

In some cases, some of the beam pairs used in the 2-D reconstructionprocess contain poor quality data, for example because of poor SNR dueto high penetration depth. In these cases, the overlap between the twobeams, which view the same depth cross section from opposite directions,may be insufficient for performing artifact compensation. When theoverlap region is not sufficient in some beam pairs, 2-D reconstructionof a slice can be carried out in two phases. First, 2-D reconstructionis performed using the beam pairs that have adequate overlap. When thebeams of a particular beam pair do not have sufficient overlap, theattenuation and time delay of this beam pair can be corrected using the2-D reconstructed values of the other beam pairs in the slice. Inapplications that can tolerate a lower refresh rate, such as non-cardiacapplications, the penetration depth (and hence the achievable overlapbetween opposite beams) can be significantly improved by using longerpulse sequences.

It should be noted that artifacts due to spatial aberrations andmulti-path are reduced in system 20, because of the use of relativelywide beams, and because each sample inherently contains informationaveraged over a large volume. Multi-path artifacts are also reduced dueto the relatively strong tissue attenuation involved.

In some embodiments, known super-resolution methods can be used inpost-processing to further improve the final spatial resolution of thesystem. Such methods may comprise, for example, Minimum Variance Methods(MVM), Burg and/or Yule-Walker methods. Super-resolution methods aredescribed, for example, by Borison et al., in “Super-Resolution Methodsfor Wideband Radar,” The Massachusetts Institute of Technology LincolnLaboratory Journal, (5:3), 1992, pages 441-461, which is incorporatedherein by reference. Super-resolution methods may be applied to eachbeam pair, to each effective slice (i.e., to the data corresponding to aspecific height of the cylinder, after 2-D reconstruction), or to thecomplete 3-D grid.

The 3-D reconstructed tissue parameters are displayed to a user by unit100 on displays 76 or 83. Different visualization methods and modes canbe used. For example, unit 100 can display isometric views of the targetorgan or parts thereof, with selected tissue parameters shown ascolor-coded layers on the 3-D display. The display can changedynamically, in accordance with the refresh rate used.

Unit 100 may display selected 2-D slices, projections and other surfacesbased of the 3-D information. Real-time 3-D data rendering can be used.Additionally or alternatively, tissue parameters can also be displayednumerically. Typically, the user can control and customize the differentdisplay functions using input devices 80. In some embodiments, unit 100enables the user to perform measurements, such as various length andvelocity measurements, based on the 3-D display.

3-D Motion Vector Analysis Mode

In the motion vector analysis mode, system 20 tracks and displays tissuedynamics, such as blood flow. The system produces real-time imagesshowing the 3-D velocity vector of each point in the scanned volume.Tissue velocity is measured using the Doppler effect.

FIG. 11 is a flow chart that schematically illustrates a method forcalculating three-dimensional motion vectors using Doppler measurements,in accordance with an embodiment of the present invention.

The method begins with system 20 scanning the target organ, at ascanning step 280. The system scans the organ from multiple angles usingpairs of opposite beams. In each beam position, several pulse sequencessimilar to the sequences of FIG. 7 above are transmitted and analyzed.The number of pulse sequences in each beam is on the order of ten,although other values can also be used.

In order to adequately image a relevant range of velocities, themeasurements should have a Nyquist frequency of on the order of 3 m/sand a velocity resolution on the order of 0.3 m/s. In order to reachthis performance, a refresh rate on the order of 20 Hz is used. In someembodiments, the beams are scanned in an interleaved manner,transmitting a single pulse sequence at a time and returning severaltimes to the same beam position, until the desired number of pulsesequences is transmitted. As a result, the pulse sequences aredistributed evenly within the overall scanning cycle. Typically, narrowbeams having azimuth and elevation beam widths on the order of 4° areused.

As described below, the Doppler shift measured with respect to a giveneffective antenna provides information related to a single component ofthe 3-D tissue velocity vector. In order to reconstruct the 3-D velocityvectors, three or more datasets are acquired. Each dataset providesinformation for each point in space, from a different viewing angle. Thedatasets can be acquired, for example, by changing the verticalinclination of the beams from one dataset to another.

For each angular beam position and for each range gate, the measurementsare de-convolved to achieve a resolution on the order of 1 mm along thevertical (cylinder height) dimension. The DSP unit then performsartifact compensation on the measurements of each beam pair, at acompensation step 284. The compensation process is similar to theprocess carried out in steps 248 and 252 of FIG. 9 above.

At the end of this process, at least three sets of measurements areavailable in cylindrical coordinates, having a spatial resolution on theorder of 1 mm. Since there is a significant overlap in azimuth betweenadjacent beams, de-convolution can also be performed between adjacenthorizontal beam angles, per range gate and height slice. De-convolutionmay also be performed between adjacent height slices, per horizontalbeam angle and range gate.

The system calculates the Doppler spectrum of each effective range gatein the scanned volume, as viewed from the direction of each beam, at aspectrum calculation step 288. Using the well-known Doppler equation,the Doppler frequency shift f_(d) of a particular tissue element can bewritten as

$f_{d} = {\frac{1}{\lambda}\left( {{{\overset{\rightarrow}{V}}_{t} \cdot \frac{{\overset{\rightarrow}{R}}_{s}}{{\overset{\rightarrow}{R}}_{s}}} + {{\overset{\rightarrow}{V}}_{t} \cdot \frac{{\overset{\rightarrow}{R}}_{r}}{\overset{\rightarrow}{R}}}} \right)}$wherein λ denotes the signal wavelength, {right arrow over (R)}_(s)denotes the position vector of the tissue element with respect to thetransmitting effective antenna, {right arrow over (R)}_(r) denotes theposition vector of the tissue element with respect to the receivingeffective antenna (which may be the same as the transmitting antenna),and {right arrow over (V)}_(t) denotes the velocity vector of the tissueelement. The Doppler spectrum of each tissue element can be calculated,for example, by applying a 16-bin discrete Fourier transform (DFT) tothe time-dependent data of each range gate in each beam position. Zeropadding may be used to improve the DFT resolution. In some embodiments,a window function may be used to reduce the level of spectral sidelobes.

In some cases, the Doppler shifts measured by a particular effectiveantenna may have different values with respect to different elements ofthe effective antenna. This effect may cause some smearing in theDoppler spectrum. The smearing effect can be estimated and compensatedfor. For example, it may be assumed that the angular Doppler velocitymeasured by a certain antenna element varies in proportion to cos(θ),wherein θ denotes the angle between the normal to the antenna plane atthe phase center of the effective antenna and the normal to the antennaplane at the location of the element in question. Using this assumption,suitable time-and-frequency dependent weighting can be applied to theDFT coefficients.

In some cases, the Doppler spectrum is also smeared as a result of thelarge signal bandwidth. In order to compensate for this effect, the DFTresult can be de-convolved with a point spread function (PSF) thatrepresents the smearing effect. The PSF can be pre-calculated using theDoppler equation, for any given spectral distribution of the transmittedsignal.

The system identifies dominant spectral components in the Dopplerspectra, at a component identification step 292. Dominant Dopplercomponents typically comprise spectral lines, or frequencies, havingrelatively strong intensities. Any suitable peak detection method knownin the art can be used to determine the peak frequencies. Centroid-basedmethods can be used to improve the calculation accuracy.

The system attempts to find dominant components that are associated withone another, i.e., relate to the same tissue velocity vector, indifferent spectra measured from different directions. Using theassociated dominant frequencies, the system solves the Doppler equationand calculates the 3-D motion vectors, such as using 3-D triangulation,at a vector calculation step 296.

In many practical cases, however, the Doppler spectra comprise multipledominant components, and it is sometimes difficult to determine whichcomponents in the different spectra are associated with one another.Calculating motion vectors based on Doppler components that are notassociated with one another usually produces false, ambiguous motionvectors.

The system detects and discards the false motion vectors, at adiscarding step 300. The system may use any suitable method or criterionfor detecting false motion vectors. For example, it may be assumed thattrue motion vectors are continuous in both time and space. In otherwords, a newly-calculated motion vector should not significantly differfrom previously-calculated vectors of the same tissue element, andmotion vectors in adjacent tissue elements should not differsignificantly from one another. Using these assumptions, motion vectorsthat vary significantly over time and/or space can be discarded. Whenusing this method, care should be taken at the boundaries of bloodvessels, whose velocity is close to zero.

The process described above produces the 3-D motion vectors of differentpoints in the scanned volume. This information can be displayed to thephysician either independently or in conjunction with the information ofother imaging modes of the system.

As noted above, the beams used in motion vector analysis are relativelynarrow. Scanning the target organ with narrow beams may limit theachievable refresh rate. In order to reduce the scanning time, in someembodiments the system transmits with a wide beam and receives withmultiple narrow beams simultaneously. The narrow beams point towards thecenter of the cylinder and have different phase center locations. Inorder to receive on multiple beams simultaneously, the receiver of thesystem should comprise multiple parallel receive chains.

The allocation of antenna elements 36 to each receive chain shouldsupport data acquisition using opposite beam pairs, in order to performartifact compensation. For this purpose, in some embodiments the receivebeams corresponding to different receive channels are configured to havedifferent phase centers, and the boresight of each receive beam isconfigured to be perpendicular to the antenna surface at thecorresponding phase center.

FIG. 12 is a block diagram that schematically illustrates a digitalreceiver and exciter unit 320, in accordance with another embodiment ofthe present invention. In unit 320, the received signal is processedsimultaneously by multiple receive chains 324. Each receive chain 324 issimilar in structure to the receiver of FIG. 5 above. The switchingarray is configured to provide each receive chain with the signal of theappropriate receive effective antenna.

Since some elements of antenna array 36 are used by multiple effectiveantennas and multiple receive chains, each of the signals received fromswitching array 92 is split, e.g., using a 1:10 splitter, and fed to adifferent combiner in each receive chain 324. Each such combiner,corresponding to a specific receive chain, may use a differentapodization scheme, so as to define the required receive beam pattern.In any given receive beam, unused elements 36 are typically given a zeroweight.

RF Therapeutic Modes

In the RF therapeutic modes (RF ablation, local heating andelectromagnetic pressure modes), system 20 applies concentrated RFenergy to a 3-D target region in the target organ. The description thatfollows refers mainly to RF ablation. Generalization to the othertherapeutic modes is straightforward.

Typically, system 20 performs RF ablation interleaved with activeimaging of the target region and its vicinity, in order to visualize andguide the ablation process. In some embodiments, the system alternatesin time between ablation frames and imaging frames. During the imagingframes, the system performs active 3-D imaging of the target organ, asdescribed above. During the ablation frames, some or all of antennaelements 36 transmit high frequency RF pulses focused on the targetablation region. The 3-D imaging information is used to adaptively trackthe focal point (i.e., the location of the target region) to whichablation energy is focused.

Adaptive tracking of the ablation region location enables ablation indynamic organs, such as the cardiac muscle. In order to enable precisetracking, the maximum frame-to-frame motion of the target region shouldgenerally not exceed a certain fraction of the ablation resolution.Frame refresh rates on the order of 50 Hz for imaging and 50 Hz forablation are typically sufficient.

Since imaging and ablation are performed using the same coordinatesystem, and since the two modes are affected by the same physicalartifacts, distortion in the acquired 3-D image typically has littleinfluence on the focal point location accuracy. Additionally, the timedelay between each element 36 and the target region, measured in theimaging mode, can be used to optimize the time delay of eachtransmitting element for ablation, so that pulses from the differentelements reach the target region simultaneously.

According to simulated results, the 3 dB width of the ablation region inapproximately 0.6 by 0.6 by 3.6 cm, when using a transmission frequencyof 18 GHz. Unlike the pulse sequences used for imaging, ablation may beperformed using relatively long pulses, so that narrow bandwidths can beused.

In some embodiments, system 20 comprises one or two additionaldome-shaped antenna arrays. The combination of the dome-shaped arraysand the cylindrical array approximate a spherical array. In a sphericalarray, the resolution is substantially uniform in all axes. Such aconfiguration is expected to provide a spatial ablation resolution ofapproximately 0.6 cm in all axes.

Additionally or alternatively, system 20 can use directional radiatingelements for ablation. The angle between the boresight of eachdirectional element and the normal to the antenna plane at the center ofthe element can be optimized to maximize the focusing of the ablationbeam, with minimal effect on the spatial resolution of the activeimaging mode. For example, all directional elements may be oriented sothat their boresights point to the center of the cylinder.

The spatial resolution of the ablation mode can also be improved byincreasing the frequency of the ablation signal. On the other hand,signal absorption in tissue increases with frequency, so that thetransmit power should be increased. It can be shown that the radius ofarray 32 has little effect on the ablation spatial resolution.

Unlike the other modes of system 20, in which processing may beperformed off-line, the processing in the RF ablation mode is performedin real-time. Assuming a volume of 5 cm², which is imaged with a spatialresolution of 2 mm per axis, the number of slices is reduced from 250 to25. Each slice comprises 360 beam positions. For each pulse, only sixrange gates are processed for each A/D converter. In order to obtain a 2mm resolution, 8 A/D converters should be used, providing 48 samples perpulse, or 96 samples for each pulse sequence (when using two pulsesequences in each transmission). For a frame rate of F frames/second,the number of samples per second is 25·360·96·F, or 8.64·10⁵·Fsamples/second. For F=50 frames/second, the acquired data rate is4.32·10⁷ samples/second.

Assuming the data rates given above, a representative digital signalprocessor (DSP) having a processing power of 15G floating-pointoperations per second (Gflop) can perform approximately 330 operationsper sample, or 33000 operations per pulse. Higher processing power canbe obtained by using multiple signal processors in parallel. Eachprocessor may process a different set of pulses, and the resulting 3-Dimage matrices can be summed, to produce the final image.

The processing load per pulse sequence is made of three dominantfactors:

(1) De-convolution per each effective range gate. There are forty-foureffective range gates per pulse sequence, corresponding to a √{squareroot over (3)}·5 cm wide region. In some cases, a √{square root over(2)}·5 cm width is also sufficient. A linear combination of eight A/Dsamples (8 additions and 8 multiplications) is calculated for each pulsesequence. The results of the two pulse sequences are added (i.e., 16operations×2 pulse sequences, plus one addition, per effective rangegate). Thus, a total of approximately 33×44=1452 operations areperformed per pulse sequence.

(2) Time delay and attenuation correction: Approximately 500 operationsare performed per pulse sequence, i.e., 1000 operations per each pair ofopposite beams.

(3) 3-D reconstruction: Each of the twenty-five samples is spread overapproximately 25×25 grid cells, producing a total of 15625 operationsper pulse sequence.

In total, the overall number of operations per pulse sequence isapproximately 20000. For F=50 frames/second, 25·360·20000·50, orapproximately 9·10⁹ operations are performed per second.

Once the 3-D image is produced, additional calculations are performedfor tissue tracking in order to adaptively track the ablation focalpoint. Tracking may be carried out by cross-correlating 3-D blocks insuccessive frames. Alternatively, tracking may be performed bydetermining the 3-D translation that minimizes the minimum mean-squareerror (or minimum mean absolute difference) between 3-D blocks insuccessive frames. For each possible motion vector, 25³ multiplicationsand additions are performed (i.e., a total of 31250 operations). Thus,1000 possible motion vectors (10 possibilities in each axis) correspondto approximately 31.25·10⁶ operations per frame. For F=50 frames/second,approximately 1.6·10⁹ operations per second are added.

In summary, a 15 Gflop DSP is typically capable of performing thecombined imaging and ablation processing in real-time. Note, however,that if the overlap between opposite beams is insufficient, theprocessing becomes more complex. In such cases, quicker, lower precisiontime delay and attenuation correction methods can sometimes be used.

Passive Imaging Mode

Different cell types in the human body, such as nerve and muscle cells,show substantial electrical activity. The propagation velocity withinneurons is typically in the range of 30-120 m/s. The duration of actionpotentials is usually in the range of 1-10 ms, and their amplituderanges between 70 and 110 mV. The duration of neuron receptor-potentialsand synaptic-potentials is typically between 5 and 100 ms, and theiramplitude ranges from 0.1 to 10 mV. These properties are described, forexample, by Kandel et al., in “Principles of Neural Science,”McGraw-Hill, New York, 2000, pages 19-35, which is incorporated hereinby reference. The propagation velocities in muscle cells tend to besignificantly lower (e.g., it takes about 50 ms for the signal generatedby the sinoatrial node in the heart to reach the atrioventricular node).

This electrical activity generates electromagnetic fields. Whenoperating in the passive imaging mode, system 20 senses and maps thesefields. The system provides a dynamic 3-D display showing thepropagation of electrical signals in a target organ, such as along anerve or a muscle. The passive and active imaging modes may be combined,so that the system displays the electrical activity overlaid on a 3-Dimage of the organ produced by active imaging.

Unlike the active imaging mode in which the signal is subject to two-waybody attenuation, in passive imaging the signal is only attenuated onits way from the organ to the receiving antenna. Furthermore, theantenna array and receiver front end may be cooled, so that theirsensitivity is improved.

In order to detect the relatively weak electromagnetic fields generatedby the physiologic electrical activity, the receiver typicallyintegrates over relatively long time intervals. For example, thereceiver may use a single A/D converter operating at a sampling ratehigher than 1 GHz, typically on the order of 10 GHz, and integrate thedata over intervals on the order of 1 μs.

In many practical scenarios, because of the low duty cycle of the sensedelectrical activity signals, the receiver will typically sense no morethan a single signal within each 1 μsec interval. Narrow receiving beamscan be used to further reduce the probability of sensing multiplesignals simultaneously.

In some embodiments, system 20 passively senses the electrical activitysignals using three or more beams having different orientations. Thebeams can be received either simultaneously or in alternation. Based onthese measurements, the system estimates the 3-D coordinates of thesource of the electrical activity, such as using interferometry methods.

Although in some cases using three beams is sufficient, a higher numberof beams (e.g., five beams) may be preferable, for example in order toreduce time delay effects caused by light velocity variations in tissue.In some embodiments, different subsets of three beams are used tocalculate a single 3-D coordinate estimate. The center of mass of thevarious estimates is used as an estimate of the signal sourcecoordinate. Alternatively, a local time delay map produced by the activeimaging mode may be used to calculate an iso-time surface, whichcorresponds to the range measured by each beam. The intersection pointof the surfaces corresponding to the three (or more) beams shouldprovide a good estimate of the signal source coordinate.

In order to perform the interferometric triangulation calculation, thesensed signal should be processed at a narrow bandwidth. In someembodiments, the sensed signal is filtered using either digital oranalog filtering.

Transmit Power Considerations and Experimental Results

The power level that should be used for transmission in system 20depends primarily on the body attenuation in the relevant imagingscenarios. In the active imaging and motion vector analysis modes, thetransmitted power should provide the system with sufficient SNR at themaximum desired penetration depth. On the other hand, the transmittedpower should not exceed industry-standard safety recommendations.

In some embodiments, the instantaneous transmitted power of the system(i.e., the instantaneous power summed over the elements of a particulareffective antenna) is set to 3 kW in the active imaging and motionvector analysis modes. As will be shown below, this power level isexpected to achieve the desired system performance, and also complieswith accepted radiation safety limits.

Using the pulse sequences described above, the transmission duty cycleis approximately 11.6%, so that the average power is approximately 350W. Assuming the body area exposed to the radiation is on the order of 1m², the average power per unit body area is 35 mW/cm². Instruction#6055.11 of the U.S. Department of Defense (DoD) entitled “Protection ofDoD Personnel from Exposure to Radiofrequency Radiation and MilitaryExempt Lasers,” Feb. 21, 1995, specifies for exposure durations shorterthan 0.1 hours a maximum permissible exposure level of 1/T_(exp) mW/cm²in any interval of T_(exp) hours. For exposure durations longer than 0.1hours, a maximum of 10 mW/cm² is permitted. Instruction #6055.11 alsostates that the restrictions on maximal exposure levels may be relaxedin cases of partial body exposure. Since imaging using the RFIT systemonly involves several seconds of exposure, the system power level iswell within the recommended safety limit.

In order to verify that the specified power level can achieve thedesired system performance, the propagation of RF energy in varioustissue types was measured using an experimental setup. The objectives ofthe experiment were (1) to quantify the expected tissue attenuation ofdifferent tissue types, and (2) to quantify the difference in tissueattenuation, reflection and time-delay values between the differenttissue types.

The experiment was conducted in an anechoic chamber. Short RF pulses inthe range 8-18 GHz were generated and analyzed by a time-domain networkanalyzer connected to a horn antenna. Different tissue samples wereirradiated with the RF pulses. Nine lamb tissue samples were tested,namely blood, heart, lungs, bone, liver, kidney, intestine, brain andthigh muscle. Two additional samples containing air and water were alsomeasured.

In each measurement, a particular sample was placed in the anechoicchamber and irradiated with RF pulses from a distance of approximately65 cm. A double-sided copper-plated PVC sheet was placed behind theirradiated sample in order to reflect the radiation back to the hornantenna and network analyzer. The area of the reflecting sheetapproximately matched the size of the sample and the beam-width of theantenna. In some of the measurements, a cascade of two different samplesplaced one behind the other was tested.

The signal generated by the network analyzer was a stepped-frequencysignal, covering the relevant frequency range in 801 pulses of differentfrequencies. Each pulse was 192 ns wide. The output power used was 1 mW.Different frequency ranges, such as 8-12 GHz, 15-18 GHz and 8-18 GHz,were tested. The network analyzer sampled the signal reflected from thesample and reflecting plate and reconstructed the signal as a functionof time. The setup was pre-calibrated to enable measurement of the netattenuation of the sample.

The network analyzer measurements provided the reflection, attenuationand time delay characteristics of each tissue type. The measured resultsclearly show that there are significant differences between differenttissue types in all three parameters. The absolute attenuation valuesmeasured support the feasibility of RFIT system 20 achieving adequateSNR at penetration depths on the order of 30 cm, enablinghigh-resolution cardiac imaging.

The experiment results were used to verify the transmit power rating ofsystem 20. Since some of the system parameters in the experiment aredifferent from the parameters the RFIT system, the calculation shouldtake these differences into account.

According to the well-known radar equation,

${{S\; N\; R} \propto \frac{P_{t}G_{a}^{2}G_{c}G_{i}}{R^{4}L}},$or, in logarithmic representation, SNR=P_(t)+2G_(a)+G_(c)+G_(i)−4R−L+C,wherein SNR denotes the signal to noise ratio of the system, P_(t)denotes the transmitted power, G_(a) denotes the antenna one-way gain,G_(c) denotes the processing gain due to pulse compression, G_(i)denotes the gain due to integration over multiple pulses, R denotes therange, and L denotes the system and medium (air and tissue) losses.

The following table shows the parameters affecting the achievable SNRboth in the experimental setup and in system 20:

RFIT Experimental Experimental RFIT system system Parameter setup[linear] setup [dB] [linear] [dB] P_(t) 10⁻³ W −30 dBW 3 · 10³ W 34.8dBW G_(a) 16 12 dB 1508 31.8 dB G_(c) 1.54 · 10⁶ 62 dB 64 18 dB G_(i)  10 dB 2 3 dB R 0.6 m −2.2 dB-m 1.7 m 2.3 dB-m L 32 15 dB X X_(dB) (mediumonly) A_(e) N/A N/A 0.075 m² −11.2 dB

Preserving the SNR of the experimental setup in the RFIT system yields:X _(dB) ={tilde over (P)} _(t)+2{tilde over (G)} _(a) +{tilde over (G)}_(c) +{tilde over (G)} _(i)−4{tilde over (R)}−(P _(t)+2G _(a) +G _(c) +G_(i)−4R−L) =34.8+63.6+18+3−9.2−(−30+24+62+0+8.8−15)=60.4 dBwherein the parameters marked ˜ denote parameters of the RFIT system andparameters not marked with ˜ denote parameters of the experimentalsetup. Practically, the SNR of the experimental setup is significantlybetter than the SNR required in system 20, so we may safely assumeX_(db)=75 dB. Based on the experimental results, such a valuecorresponds to a penetration depth of approximately 30 cm. Thispenetration depth is suitable for high penetration depth applications,such as cardiac imaging.

In the RF ablation mode, the transmitted power of system 20 shouldenable raising the temperature of the target ablation region byapproximately 20° C. As described above, the dimensions of the ablationregion are approximately 0.6 by 0.6 by 3.6 cm, at the 3 dB points.Assuming the ablation region has the shape of an ellipsoid, its volumeis approximately (4π/3)·0.3·0.3·1.8=0.68 cm³. Assuming a characteristictissue density of 1 g/cm³, the mass of the ablation region isapproximately 0.7 grams. The energy required to increase the temperatureof this mass by 1° C. is 2.9 Joules. Thus, approximately 59 Joules arerequired for increasing the temperature of the ablation region by 20° C.

Assuming the ablation procedure is 600 second long, the power dissipatedin the ablation region should be somewhat greater than 59/600=0.098 W.After reaching the target temperature, the power used is normallydecreased, so that the local temperature remains stable (i.e., so thatthe body heat dissipation mechanism and the RF heating mechanism balanceeach other).

According to the results of the experiment described above, the maximumone-way power attenuation is approximately 2 dB/cm. This attenuation wasmeasured for blood and water. A pulse traveling a distance of 20 cm intissue would therefore be attenuated by no more than 40 dB. According tocomputer simulation results, the ratio between the transmitted power perelement and the power at the ablation region is 1:27712 forcontinuous-wave transmission. Thus, the transmission power per element Pshould be set to approximately:

$P = {{\frac{0.098 \cdot 10000}{27712}W} = {35.4\mspace{14mu}{mW}}}$

In the simulated scenario, an array comprising 29964 elements is used,so that the overall RMS power is approximately 1060 W. Assuming theirradiated body area is 1.5 m², the power per unit area is 70 mW/cm²,which is well within the short-exposure limits of the DoD safetystandard cited above. For longer durations, on the order of 600 seconds(0.1 hours), the power density is acceptable for treatment scenarios.

The calculation above is based on worst-case attenuation assumptions,and in many practical cases the transmitted power can be significantlylower. When system 20 uses additional dome-shaped antenna arrays, thedimensions of the ablation region can be reduced to approximately 0.6 by0.6 by 0.6 cm. In such configurations, the energy required for ablationis reduced by a factor of six, and the power per unit body area isreduced to approximately 11 mW/cm². In some embodiments, sensitive bodyparts can be covered by radiation absorbing materials in order tofurther reduce their exposure. It should be noted that the heatingeffect decays rapidly outside the 3 dB ablation region. Typically,immediately outside the 3 dB boundary, each tissue point is heated by0.5° C. or less for each 1° C. temperature increase inside the ablationregion.

The experimental setup described above was also used to verify thefeasibility of enhancing the range resolution using multiple A/Dconverters. The reflecting plate and tissue sample were placed atmultiple distances from the antenna. The distances were spaced byapproximately 1.57 mm from one another, over a 15.7 mm range. Thedistance offsets are equivalent to the time delay offsets between A/Dconverters in system 20.

The data measured by the network analyzer was processed in off-line toextract the signal magnitude per each effective range gate. The resultsclearly showed that the measured signal changed significantly when thesample was moved by increments smaller than the raw range resolution.The measured ranges to the different reflecting surfaces of the sampleand plate approximately matched the manually-measured distances.

System Test and Calibration

In some embodiments, different test and calibration procedures may beperformed at different life cycle stages of system 20. For example,during installation at a particular site, the installation quality andsystem integrity are evaluated and corrected if necessary. Attention istypically given to mechanical deformations in array 32, which may affectthe relative locations of the various radiating elements. The cablelengths, the accuracy of delay-lines and the attenuation of cables andconnectors, which may affect the beam-shaping of the various effectiveantennas used, are also sometimes tested at installation.

During everyday use of the system, the system may be tested both at theradiating element level and at the effective antenna level. Such testsmay be performed periodically or when a fault is detected or suspected.

Radiating element level calibration and test may be performed using anaccurately-manufactured calibration sphere, which is positioned at thecenter of the cylindrical antenna array. For each element 36, pulses aretransmitted and received using the element. Receiving a signal whosepower is within an expected range is indicative of the integrity of theentire transmit-receive chain. The power of the received signal may alsobe used for compensating for the attenuation along the two-way path fromthe exciter to the element. Furthermore, the minimal range at which anon-negligible reflection is detected, corresponding to the distance tothe sphere surface, can be used for minimizing range bias. Othercalibration procedures may involve transmitting pulses from a particularelement 36, and receiving the signal at several adjacent elements. Therelative timing of the reflections can be used for minimizing phasemisalignment between adjacent elements.

At the effective antenna level, a specially-manufactured phantom (i.e.,an artificial body imitating an imaged object) can be placed atpre-determined locations within the cylindrical array. Pulses aretransmitted from each possible effective antenna. The reflection fromthe phantom is received by the effective antenna, as a function ofrange. Separate measurements can be performed for each frequencysub-band. The minimal range at which a non-negligible reflection isdetected can be used for range calibration. Significant deviations fromthe expected signal may indicate failure or performance degradation.

In some embodiments, as described above, the background signal (i.e.,the signal output by the receiver without the presence of a reflectingobject) is subtracted from the output of each received beam during imagereconstruction. The background signal is typically pre-measured for eachpossible effective antenna. The background signal of each effectiveantenna is often measured as a function of range and for each frequencysub-band.

Data Acquisition Load Estimation

As noted above, measured data can be collected in system 20 either inreal-time or off line (with the exception of RF therapeutic mode, inwhich data is typically collected on-line). When operating off-line, theacquired data is stored in DSP unit 96 or in data processor unit 100.Typically, the duration of a complete off-line data acquisition cycle ison the order of 1 second.

In the active imaging mode, fifteen A/D converters sample the signalsimultaneously. Assuming a 10 GHz signal bandwidth, a range gate of 1.5cm and a region of interest that is 50 cm deep, the number of samplesper pulse per A/D converter is 35. When using B bits per sample (e.g.,B=20 bits, 10 bits per complex component), the number of bits per pulseper A/D converter is 35·B. For 15 simultaneous A/D converters, thenumber of bits per pulse is 525·B. For a pulse repetition frequency(PRF) of 18 MHz, the data acquisition rate is 9.45·B Gbit/sec.

In the motion vector analysis mode, in comparison to the active imagingmode, datasets are collected from three different directions, at a lowerrefresh rate (typically one fifth of the refresh rate used for activeimaging). Multiple pulse sequences, typically 10 sequences, are used perbeam position. The data acquisition rate is thus 3·10/5=6 times higherthan the data rate of the active imaging mode, or 56.7·B Gbit/sec.

In the RF therapeutic modes, data acquisition is performed for real-timeimaging and guidance. The frame rate used is typically half the rateused in active imaging, and the number of vertical slices used istypically one tenth of the number of slices used in active imaging.Moreover, a smaller region of only 5 cm³ region is typically imaged. Thenumber of samples per pulse is approximately 3.8 times lower than inactive imaging, yielding a data acquisition rate of approximately 125·BMbit/sec.

In the passive imaging mode data is collected at a very high rate,typically on the order of 10 GHz. Data is collected from M beamssimultaneously. Each sample uses B bits. The resulting data acquisitionrate is 10·M·B Gbit/sec.

Although the embodiments described herein mainly address cardiacapplications, the methods and systems described herein can also be usedfor imaging and applying treatment to any other organ or system. RFITmethods and systems can be used for detecting and removing tumors, suchas in the different fields of oncology. The tissue classificationcapability can be used for tissue analysis in pathology. Measuring localconductivity can be used in different neurological applications. Otherapplications may include veterinary applications and general clinicalresearch.

Variations of the RFIT system can be used as all-in-one diagnosis and/ortreatment tools in environments having limited access to hospitals, suchas remote rural areas, oil barges, ships and space stations.

RFIT systems can apply additional types of treatment. For example,assuming plaque or other material deposited in arteries is particularlysensitive to heat in comparison with the surrounding tissue, a low powerRFIT system can use RF ablation to treat arterial stenosis, such as inthe coronary arteries. RF ablation can also be used to destroy emboli.An additional application of the local heating mode may be theactivation of temperature-activated drugs. Such drugs can be introducedinto the patient body and activated only in a particular location ororgan by applying local heating.

Other applications of RFIT methods and systems may be in orthopedics andsports medicine. For example, the capability to perform 3-D imaging atextremely high refresh rates enables gathering kinetic data regardingskeletal motion and the motion of muscles.

The different RFIT modes, and in particular the local heating mode, canbe used in different para-medical applications. For example, localheating can be used in cosmetics, such as for treating differentdermatological conditions. Applications may also exist in alternativemedicine.

RFIT methods and systems can also be used in non-medical applications.For example, an RFIT system can be used for 3-D modeling oftemperature-sensitive materials and 3-D imaging of non-metallic objectsin various industrial applications.

RFIT methods and systems can also be used for security applications. Forexample, the ability to measure conductivity with high spatialresolution can be used to remotely detect concealed weapons. A variationof the local heating mode can be used to temporarily incapacitate aperson identified as carrying a concealed weapon. The tissuecharacterization capability of the system can be used to detectexplosives, drugs and other illegal substances.

The RFIT systems described hereinabove are based on backscatteringreflections. An alternative high resolution imaging system may also bebased on attenuation. The attenuation-based system can use a cylindricalarray of radiating elements, which transmits relatively wide beams,approximately corresponding to the width of the imaged organ. Theattenuated signal is received by multiple narrow beams concurrently. Thereceive beams may either span several locations in azimuth, or severallocations in both azimuth and elevation. Both transmit and receive beamsshould point to the long axis of the subject, but receive and transmitbeams should be located at different sides of the subject. The transmitbeams may span 180° or 360° around the circumference of the cylinder.

In principle, an attenuation-based system can provide a two-dimensionalor one-dimensional array of attenuation parameters for each transmitbeam position. Improved performance may be obtained by transmitting atseveral elevations along the cylinder height. Image reconstruction canuse known methods, such as reconstruction methods used in computerizedtomography (CT) imaging systems.

In order to reduce the effects of refraction, which cause the beams todeviate from a straight line, each receive beam may be relatively wide.In such cases, adjacent receive beams should have an overlappingmainlobe, and the angular resolution can be achieved by de-convolutionprocedures.

Attenuation-based systems can use narrowband RF signals, so as tosimplify the hardware used. In addition, the power used should onlyassume one-way attenuation, so that high penetration depths may beachieved. In some embodiments, the attenuation-based system can usemechanical antenna scanning. A wide-angle source (e.g., a horn antenna)may be placed on one side of the subject, and an array of multiplereceive beams (e.g., an antenna with digital beam-forming or an array ofdiscrete antennas) can be placed on the opposite side. The two antennasare mechanically-scanned around the subject in one or two dimensions. Inapplications in which a high refresh rate is not mandatory, asingle-beam receiving antenna may be used, and the multiple receivebeams can be generated by transmitting a series of pulses for each beamposition, and moving the receive antenna from pulse to pulse.

It will thus be appreciated that the embodiments described above arecited by way of example, and that the present invention is not limitedto what has been particularly shown and described hereinabove. Rather,the scope of the present invention includes both combinations andsub-combinations of the various features described hereinabove, as wellas variations and modifications thereof which would occur to personsskilled in the art upon reading the foregoing description and which arenot disclosed in the prior art.

1. A method for imaging, comprising: directing a plurality of radio frequency (RF) beams toward a target organ from a respective plurality of antenna locations, the plurality of the RF beams comprising one or more first pairs of the RF beams, each pair comprising two of the RF beams that impinge on the target organ from opposite directions; receiving RF signals reflected from the target organ responsively to the RF beams, the RF signals comprising one or more second pairs of the RF signals engendered respectively by the one or more first pairs of the RF beams; compensating for local tissue artifacts in the RF signals by jointly processing the RF signals in each of the second pairs; and using a processor to calculate three-dimensional (3-D) velocity vectors of multiple points in the target organ with respect to the antenna locations using the RF signals after compensating for the local tissue artifacts.
 2. The method according to claim 1, wherein calculating the 3-D velocity vectors comprises evaluating Doppler spectra of the RF signals with respect to the antenna locations for each of the multiple points, identifying dominant spectral components in the Doppler spectra and associating three or more of the dominant spectral components in respective three or more of the Doppler spectra to produce a 3-D velocity vector estimate.
 3. The method according to claim 2, wherein associating the three or more dominant spectral components comprises identifying and discarding false associations between dominant spectral components by comparing the 3-D velocity vector estimate to at least one estimate selected from a group of estimates consisting of previous 3-D velocity vector estimates and 3-D velocity vector estimates of adjacent points in the target organ.
 4. An imaging system, comprising: a set of antennas, which are arranged to direct a plurality of radio frequency (RF) beams toward a target organ from a respective plurality of antenna locations, the plurality of the RF beams comprising one or more first pairs of the RF beams, each first pair comprising two of the RF beams that impinge on the target organ from opposite directions; a receiver, which is arranged to receive via the set of antennas RF signals reflected from the target organ responsively to the RF beams, the RF signals comprising one or more second pairs of the RF signals engendered respectively by the one or more first pairs of the RF beams; and a processor, which is arranged to compensate for local tissue artifacts in the RF signals by jointly processing the RF signals in each of the second pairs, and to calculate three-dimensional (3-D) velocity vectors of multiple points in the target organ with respect to the antenna locations using the RF signals after compensating for the local tissue artifacts.
 5. The system according to claim 4, wherein the processor is arranged to calculate the 3-D velocity vectors by evaluating Doppler spectra of the RF signals with respect to the antenna locations for each of the multiple points, identifying dominant spectral components in the Doppler spectra and associating three or more of the dominant spectral components in respective three or more of the Doppler spectra to produce a 3-D velocity vector estimate.
 6. The system according to claim 5, wherein the processor is arranged to identify and discard false associations between dominant spectral components by comparing the 3-D velocity vector estimate to at least one estimate selected from a group consisting of previous 3-D velocity vector estimates and 3-D velocity vector estimates of adjacent points in the target organ.
 7. The system according to claim 5, wherein the processor is arranged to estimate tissue motion velocities at the multiple points in the target organ by measuring Doppler spectra of the RF signals in three or more of the RF beams.
 8. The system according to claim 5, wherein the set of antennas is a cylindrical antenna array surrounding said target organ, and wherein the RF beams are parallel, with an offset no greater than one degree, to a base of the cylinder and point toward a central axis of the cylinder from multiple azimuth angles and heights.
 9. The system according to claim 8, wherein the cylindrical array is tilted at an angle with respect to ground.
 10. A computer-readable storage medium, having computer executable instructions, which instructions, when executed by a computer, cause the computer to control one or more antennas to direct a plurality of radio frequency (RF) beams toward a target organ from a respective plurality of angles, the plurality of the RF beams comprising one or more first pairs of the RF beams, each first pair comprising two of the RF beams that impinge on the target organ from opposite directions, to receive via the one or more antennas RF signals reflected from the target organ responsively to the RF beams, the RF signals comprising one or more second pairs of the RF signals engendered respectively by the one or more first pairs of the RF beams, to extract local tissue parameters at multiple points in the target organ by jointly processing the RF signals in each of the second pairs and to produce images of the target organ using the extracted local tissue parameters.
 11. A computer-readable storage medium, storing computer readable instructions, which instructions, when executed by a computer, cause the computer to control a set of antennas to direct a plurality of radio frequency (RF) beams toward a target organ from a respective plurality of antenna locations, the plurality of the RF beams comprising one or more first pairs of the RF beams, each first pair comprising two of the RF beams that impinge on the target organ from opposite directions, to receive via the set of antennas RF signals reflected from the target organ responsively to the RF beams, the RF signals comprising one or more second pairs of the RF signals engendered respectively by the one or more first pairs of the RF beams, to compensate for local tissue artifacts in the RF signals by jointly processing the RF signals in each of the second pairs, and to calculate three-dimensional (3-D) velocity vectors of multiple points in the target organ with respect to the antenna locations using the RF signals after compensating for the local tissue artifacts.
 12. A method for radio frequency (RF) ablation, comprising: directing a plurality of RF beams toward a target organ from a respective plurality of angles, the plurality of the RF beams comprising one or more first pairs of the RF beams, each first pair comprising two of the RF beams that impinge on the target organ from opposite directions; receiving RF signals reflected from the target organ responsively to the RF beams, the RF signals comprising one or more second pairs of the RF signals engendered respectively by the one or more first pairs of the RF beams; using a processor to extract local tissue parameters at multiple points in the target organ by jointly processing the RF signals in each of the second pairs; and focusing an ablating signal on an ablation region in the target organ using multiple ablation beams based on the extracted local tissue parameters.
 13. A radio frequency (RF) ablation system, comprising: one or more antennas, which are arranged to direct a plurality of RF beams toward a target organ from a respective plurality of angles, the plurality of the RF beams comprising one or more first pairs of the RF beams, each first pair comprising two of the RF beams that impinge on the target organ from opposite directions; a receiver, which is arranged to receive via the one or more antennas RF signals reflected from the target organ responsively to the RF beams, the RF signals comprising one or more second pairs of the RF signals engendered respectively by the one or more first pairs of the RF beams; a transmitter, which is arranged to transmit an ablating signal toward an ablation region in the target organ via the one or more antennas; and a processor, which is arranged to extract local tissue parameters at multiple points in the target organ by jointly processing the RF signals in each of the second pairs, and to cause the ablating signal to be focused on the ablation region in the target organ based on the extracted local tissue parameters.
 14. A computer readable storage medium having computer-executable instructions to control performance of the imaging method of claim
 1. 15. An imaging method comprising: (a) providing an apparatus having: at least one antenna, adapted to direct a plurality of radio frequency (RF) beams toward a target organ from a respective plurality of antenna locations, an RF receiver, adapted to receive signals reflected from the target organ responsively to the RF beams, and a processor; and (b) carrying out the method of claim 1 with said apparatus.
 16. The method according to claim 15, wherein calculating the 3-D velocity vectors comprises evaluating Doppler spectra of the RF signals with respect to the antenna locations for each of the multiple points, identifying dominant spectral components in the Doppler spectra and associating three or more of the dominant spectral components in respective three or more of the Doppler spectra to produce a 3-D velocity vector estimate.
 17. The method according to claim 16, wherein associating the three or more dominant spectral components comprises identifying and discarding false associations between dominant spectral components by comparing the 3-D velocity vector estimate to at least one estimate selected from a group of estimates consisting of previous 3-D velocity vector estimates and 3-D velocity vector estimates of adjacent points in the target organ.
 18. A computer readable storage medium having computer-executable instructions to control performance of the RF-ablation method of claim
 12. 19. An ablation method comprising: (a) providing an apparatus having: at least one antenna, adapted to direct a plurality of radio frequency (RF) beams toward a target organ from a respective plurality of antenna locations, an RF receiver, adapted to receive signals reflected from the target organ responsively to the RF beams, a processor; and a transmitter, arranged to transmit an ablating signal towards an ablation region in the target organ; and (b) carrying out the method of claim 12 with said apparatus. 